Microfluidic Liposome Synthesis, Purification and Active Drug Loading

ABSTRACT

Microfluidic methods and systems are provided for continuous flow synthesis and active loading of liposomes, which include a liposome formation region configured to form a population of liposomes and a microdialysis region downstream from the liposome formation region and configured to form a transmembrane gradient for active drug loading of the liposomes. Microfluidic methods and systems for high throughput production of liposomes are also provided featuring high aspect ratio microchannels.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a divisional of U.S. patent application Ser. No.14/524,797, filed Oct. 27, 2014, which application is based on U.S.Provisional Patent Application Ser. No. 61/896,204, filed Oct. 28, 2013,which applications are incorporated herein by reference in theirentireties and to which priority is claimed.

STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT

This work was supported by the National Science Foundation (NSF) underCBET0966407 grant. The US government has certain rights in thisinvention.

FIELD OF THE INVENTION

The present invention relates to microfluidic methods and systems forcontinuous flow synthesis of nanoparticles, and in particularmicrofluidic systems for preparing drug-loaded liposomes as well asmicrofluidic systems for high throughput production of liposomes.

BACKGROUND OF THE INVENTION

Liposomes are lipid bilayer vesicle nanoparticles which have attractedgreat interest as drug delivery vehicles (Kikuchi, H. et al. (1999)“Gene delivery using liposome technology,” J. Conto. Rel. 62:269-277;Templeton et al. (1997) “Improved DNA: Liposome complexes for increasedsystemic delivery and gene expression,” Nat. Biotechnol. 15:647-652;Abraham. S. et al. (2005) “The liposomal formulation of doxorubicin,”Methods Enzymol. 391:71-97; Andresen, T. L. et al. (2005) “Advancedstrategies in liposomal cancer therapy: Problems and prospects of activeand tumor specific drug release,” Prog. Lipid Res. 44:68-97;Ramachandran, S. et al. (2006) “Nanoliposomes for Cancer Therapy: AFMand Fluorescence imaging of Cisplatin Encapsulation, Stability, CellularUptake and Toxicity,” Langmuir 22:8156-8162; Zamboni, W. C. (2005)“Liposomal Nanoparticle, and Conjugated Formulations of AnticancerAgents,” Clin. Cancer Res. 11:8230-8234). Encompassing the ability toencapsulate aqueous solutions within their core, isolate lipophiliccompounds within their lipid bilayer, and support tailored surfacechemistries for targeted delivery, liposomes are versatile,multifunctional nanoparticles with numerous applications including drugdelivery for cancer treatment, antibiotics, and anesthetic compounds(Allen, T. M. & Cullis, P. R. (2013) “Liposomal drug delivery systems:From concept to clinical applications,” Adv. Drug Delivery Rev.65:36-48; Fenske, D. B. et al. (2008) “Liposomal nanomedicines: anemerging field,” Toxicologic Pathology 36:21-29).

Due to their ability to increase the drug loading capacity by an orderof magnitude or more compared to other delivery methods (Gu, F. X. etal. (2007) “Targeted nanoparticles for cancer therapy,” Nano Today2:14-21) and protect their payloads from metabolic activity and earlyexcretion, liposomal drugs provide increased therapeutic indices whileminimalizing the damaging side effects of their non-encapsulatedcounterparts (Immordino, M. L. et al. (2006) “Stealth liposomes: reviewof the basic science, rationale, and clinical applications, existing andpotential,” International J. Nanomedicine, 1(3):297-315). As such,liposome encapsulated drugs have had an immense impact in oncology, withlong circulating liposomes providing preferential extravasation fromtumor vessels (Park, J. W. et al. (2002) “Future directions of liposome-and immunoliposome-based cancer therapeutics,” Semin. Oncol.31:196-205).

Liposome-encapsulated drugs have exhibited potent activity against awide range of cancers including breast, ovarian, uterine, and othersolid tumors (Gu, F. X. et al. (2007) “Targeted nanoparticles for cancertherapy,” Nano Today 2:14-21; Patri, A. K. et al. (2002) “Dendriticpolymer macromolecular carriers for drug delivery,” Current Opinion inChemical Biology 6:466-471; Wang, A. Z. et al. (2012) “NanoparticleDelivery of Cancer Drugs,” Annu. Rev. Med. 63:185-98; Brannon-Peppas, L.& Blanchette, J. O. (2012) “Nanoparticle and targeted systems for cancertherapy,” Adv. Drug Deliv. Rev. 64:206-212; Gabizon, A. & Martin, F.(1997) “Polyethylene glycol-coated (pegylated) liposomal doxorubicin.Rationale for use in solid tumors.” Drugs 54:15-21; Krishna, R. & Mayer,L. D. (1997) “Liposomal doxorubicin circumvents PSC 833-free druginteractions, resulting in effective therapy of multidrug resistantsolid tumors,” Cancer Res. 57:5246-53). Liposome delivery systems havealso had an impact in vaccinology (Glück, R. (1995) “Liposomalpresentation of antigens for human vaccines,” Pharm. Biotechnol.6:325-45), ophthalmology (Ebrahim, S. et al. (2005) “Applications ofliposomes in ophthalmology,” Surv. Ophthalmol. 50:167-82), pulmonology(Schreier, H. et al. (1993) “Pulmonary delivery of liposomes,” J.Control. Release 24:209-223), and numerous other pathologies (Moghimi,S. M. & Szebeni, J. (2003) “Stealth liposomes and long circulatingnanoparticles: critical issues in pharmacokinetics, opsonization andprotein-binding properties,” Prog. Lipid Res. 42:463-478; Tazina, E. V.et al. (2011) “Qualitative and quantitative analysis of thermosensitiveliposomes loaded with doxorubicin,” Pharm. Chem. J. 46:54-59).

Tumor microvasculature is porous, with pore diameters large enough toallow nanoparticles smaller than several hundred nanometers to migrateinto the extravascular space, providing a mechanism forliposome-encapsulated drugs to concentrate within tumors (Abraham. S. etal. (2005) “The liposomal formulation of doxorubicin,” Methods Enzymol.391:71-97). Using this feature, liposomal anthracyclines have shownreduced toxicity compared with conventional delivery methods, whileproviding efficacies comparable with their conventional counterparts(O'Shaughnessy, J. (2003) “Liposomal Anthracyclines for Breast Cancer:Overview,” Oncologist 8:1-2). For example, Doxil®, the firstFDA-approved nanoparticle drug, comprising the anthracycline antibioticdoxorubicin in PEGylated ˜100 nm liposomes (Barenholz, Y. C. (2012)“Doxil®—The First FDA-Approved Nano-Drug: Lessons Learned,” J.Controlled Release 160:117-134), is widely used as a chemotherapeuticfor treatment of a range of recurrent cancers, and there are variousother liposomal drugs approved for clinical use (Chang, H. I. & Yeh, M.K. (2012) “Clinical development or liposome-based drugs: formulation,characterization, and therapeutic efficacy,” International J. ofNanomedicine 7:49-60; Wagner, V. et al. (2006) Nat. Biotechnol.24:1211-7), with many more formulations in clinical trials.

Vesicle size and polydispersity are key parameters impacting thetherapeutic index of liposomal drugs. Smaller liposomes are known toexhibit slower blood clearance rates, thereby increasing drugbioavailability (Chang, H. I. & Yeh, M. K. (2012) “Clinical developmentor liposome-based drugs: formulation, characterization, and therapeuticefficacy,” International J. of Nanomedicine 7:49-60; Litzinger, D. C. etal. (1994) “Effect of liposome size on the circulation time andintraorgan distribution of amphipathic poly(ethylene glycol)-containingliposomes,” Biochim. Biophys. Acta 1190:99-107). Various attempts havebeen made to investigate the impact of liposome size on cell uptake,intracellular transport and fate, and overall biodistribution forvesicles in the ˜100 to ˜1,000 nm range (Ramachandran, S. et al. (2006)“Nanoliposomes for Cancer Thereapy: AFM and Fluorescence imaging ofCisplatin Encapsulation, Stability, Cellular Uptake and Toxicity,”Langmuir 22:8156-8162; Kelly, C. et al. (2011) “Targeted Liposomal DrugDelivery to Monocytes and Macrophages,” J. Drug Delivery 2011, 727241;Ahsan, E. et al. (2002) “Targeting to macrophages: role ofphysiochemical properties of particulate carriers—liposomes andmicrospheres—on the phagocytosis by macrophages,” J. Controlled Release79:29-40; Epstein-Barash, H. et al. (2010) “Physicochemical parametersaffecting liposomal bisphosphonates bioactivity for restenosis therapy:internalization, cell inhibition, activation of cytokines andcomplement, and mechanism or cell death,” J. Controlled Release146:182-195; Takano, S. et al. (2003) “Physicochemical properties ofliposomes affecting apoptosis induced by cationic liposomes inmacrophages,” Pharmaceutical Research 20:962-968; Pollock, S. et al.(2010) “Uptake and trafficking of liposomes to the endoplasmicreticulum,” The FASEB 24:1866-1878). In some studies, liposomes largerthan 300 nm were not effectively taken up by cells in vitro, whilesmaller 100 nm liposomes exhibited rapid endocytosis (Ramachandran, S.et al. (2006) “Nanoliposomes for Cancer Therapy: AFM and Fluorescenceimaging of Cisplatin Encapsulation, Stability, Cellular Uptake andToxicity,” Langmuir 22:8156-8162). In other studies, 100 nm liposomeswere found to maximize uptake into monocytes and macrophages comparedwith larger vesicles (Kelly, C. et al. (2011) “Targeted Liposomal DrugDelivery to Monocytes and Macrophages,” J. Drug Delivery 2011, 727241;Ahsan, E. et al. (2002) “Targeting to macrophages: role ofphysicochemical properties of particulate carriers—liposomes andmicrospheres—on the phagocytosis by macrophages,” J. Controlled Release79:29-40; Epstein-Barash, H. et al. (2010) “Physicochemical parametersaffecting liposomal bisphosphonates bioactivity for restenosis therapy:internalization, cell inhibition, activation of cytokines andcomplement, and mechanism or cell death,” J. Controlled Release146:182-195; Takano, S. et al. (2003) “Physicochemical properties ofliposomes affecting apoptosis induced by cationic liposomes inmacrophages,” Pharmaceutical Research 20:962-968). In the case ofliposomal doxorubicin, higher tumor uptake has been observed togetherwith significantly lower uptake by healthy tissues when the mean vesiclesize was reduced from 100 nm to 75 nm (Cuia, J. et al. (2007) “Directcomparison of two pegylated liposomal doxorubicin formulations: is AUCpredictive for toxicity and efficacy?” J. Controlled Release118:204-215). As such, questions remain about the impact of liposomesize on cell uptake and internalization for vesicles smaller than about100 nm, in part due to limitations imposed by most existing liposomepreparation techniques.

Thus, a challenge to liposomal drug delivery technologies has been theeffective control over vesicle size during synthesis. The ideal size forcancer-targeting liposomal nanomedicines is commonly believed to beabout 100 nm, which is thought to be large enough to provide a high drugpayload volume while small enough to pass through leaky endothelialjunctions in tumor tissues (Fenske, D. B. et al. (2008) “Liposomalnanomedicines: an emerging field,” Toxicologic Pathology 36:21-29).However, such view ignores the increasing use of liposomes forlipophilic drug encapsulation within the vesicle membrane where loadingefficiency scales inversely with liposome size, and also ignores theimpact of vesicle size on key parameters affecting drug efficacy andsafety, including cellular uptake, cellular fate, and overallbiodistribution. Relationships between such key characteristics andliposome size are not fully understood for nanoparticles below 100 nm,in large part because many conventional bulk synthesis techniques yieldrelatively large and polydisperse liposome populations that renderdetailed size-dependent studies difficult.

Another challenge hampering liposomal delivery systems has been thedevelopment of effective methods for loading high concentrations oftherapeutic agent(s) or drug into lipid vesicles. Increaseddrug-to-lipid ratio (D/L) is a highly desirable attribute for liposomedelivery systems, given in vivo toxicity is inversely related to D/L(Mayer, L. D. et al. (1989) “Influence of Vesicle Size, LipidComposition, and Drug-to-Lipid Ratio on the Biological Activity ofLiposomal Doxorubicin in Mice” Cancer Res. 49:5922-5930). For example,nanoparticle delivery systems can enhance the therapeutic index ofanti-cancer agents by increasing drug concentration in tumor cells. Theincreased drug concentration facilitated by nanoparticle delivery is theresult of enhanced nanoparticle permeability and retention in tumortissues (Matsumura, Y. & Maeda. H. (1986) “A new concept formacromolecular therapeutics in cancer chemotherapy: mechanism oftumoritropic accumulation of proteins and the antitumor agent SMANCS,”Cancer Res 6:193-210; Maeda, H. (2010) “Tumor-selective delivery ofmacromolecular drugs via the EPR effect: background and futureprospects,” Bioconjugate Chemistry 21:797-802), together with the use ofmolecular targeting strategies that enhance tumor cell uptake (Moses, M.A., Brem. H. & Langer. R. (2003) “Advancing the field of drug delivery:taking aim at cancer,” Cancer Cell 4:337-341; Liu, Y. et al. (2007)“Nanomedicine for drug delivery and imaging: a promising avenue forcancer therapy and diagnosis using targeted functional nanoparticles,”International J. Cancer 120:2527-2537; Cho, K. et al. (2008)“Therapeutic nanoparticles for drug delivery in cancer,” Clin. CancerResearch 14:1310-1316).

A variety of liposomal drug synthesis techniques have been reported(Otake. K. et al. (2006) “Preparation of liposomes using an improvedsupercritical reverse phase evaporation method,” Langmuir The ACSJournal of Surfaces and Colloids 22:2543-2550; Uhumwangho, M. U. & Okor,R. S. (2005) “Current trends in the production and biomedicalapplications of liposomes: a review,” Sciences New York 4:9-21; Jiskoot,W. et al. (1986) “Preparation of liposomes via detergent removal frommixed micelles by dilution. The effect of bilayer composition andprocess parameters on liposome characteristics,” Pharmaceutisch WeekbladScientific Edition 8:259-265; Szoka, F. & Papahadjopoulos, D. (1978)“Procedure for preparation of liposomes with large internal aqueousspace and high capture by reverse phase evaporation,” Proceedings of theNational Academy of Sciences of the United States of America75:4194-4198; Meure, I. A. et al. (2008) “Conventional and Dense GasTechniques for the Production of Liposomes: A Review,” Aaps Pharmscitech9:798-809).

Conventional liposome preparation is based on demanding bulk-scaleprocesses which include a variety of traditional methods includingethanol injection, reverse-phase evaporation, detergent depletion,emulsification, supercritical phase formation, membrane extrusion,thin-film hydration, rapid solvent exchange, which all requirepost-processing steps such as sonication or membrane extrusion toregulate the size and reduce the polydispersity of the final populationof liposomes (Jesorka, A. & Orwar, 0. (2008) “Liposomes: technologiesand analytical applications” Annu. Rev. Anal. Chem. (Palo Alto. Calif.1:801-32). In addition, such techniques require further processing stepsfor drug encapsulation, membrane functionalization, purification, andconcentration. Conventional bulk methods are therefore cumbersome, timeconsuming, and labor intensive, and result in liposomal nanomedicineswith limited shelf life due to drug leakage and lipid degradation.Moreover, even after repeated processing steps utilizing bulktechniques, such as sequential membrane extrusion or size exclusionchromatography, the resulting vesicles tend to be relatively large (>100nm) and polydisperse. For example, when performing 6-step membraneextrusion to reduce liposome size and polydispersity, relative standarddeviations above 50% are observed (Berger, N. et al. 2001) “Filterextrusion of liposomes using different devices: comparison of liposomesize, encapsulation efficiency, and process characteristics,”International Journal of Pharmaceutics 223:55-68). In addition,conventional bulk synthesis and encapsulation methods often lead tosignificant agent loss with waste that can approach 98% for hydrophilicdrug compounds (Lasic, D. D. (1998) “Novel applications of liposomes,”Trends in Biotechnology 16:307-321; Nagayasu, A. et al. 1999) “The sizeof liposomes: a factor which affects their targeting efficiency totumors and therapeutic activity or liposomal antitumor drugs,” AdvancedDrug Delivery Reviews 40:75-87).

Because conventional bulk synthesis yields relatively large andpolydisperse liposomes, detailed size-dependent behaviors for smallervesicles have proven difficult or impossible to study. As a result,studies of size-dependent behaviors have largely focused on the use ofinorganic nanoparticles such as gold (Chithrani, B. D. et al. (2006)“Size and Shape Dependence of Nanoparticles on Cellular Uptake,” Nano668:662-668; Shan, Y. et al. (2009) “Size-dependent endocytosis ofsingle gold nanoparticles,” Chemical Communications 47:8091-8093; Zhang,S. et al. (2009) “Size-Dependent Endocytosis of Nanoparticles,” Advancedmaterials Deerfield Beach Fla. 21:419-424; Cho, E. C. et al. (2011)“Cellular uptake of gold nanoparticles,” Cancer Cell 6:385-391), carbon(Jin, H. et al. (2009) “Size-dependent cellular uptake and expulsion ofsingle-walled carbon nanotubes: single particle tracking and a genericuptake model for nanoparticles,” ACS nano 3:149-158), iron oxide (Huang,J. et al. (2010) “Effects of nanoparticle size on cellular uptake andliver MRI with polyvinylpyrrolidone-coated iron oxide nanoparticles,”ACS nano 4:7151-7160), and silica (Kumar, S. et al. (2012) “SizeDependent Interaction of Silica Nanoparticles with Different Surfactantsin Aqueous Solution,” Langmuir 28(25):9288-9297, which materials can besynthesized with relatively tight control over size and with lowpolydispersity. However, because surface properties of such inorganicnanoparticles are entirely different from liposomes, which closely mimicthe native cell walls across which endocytosis occurs, the results ofsuch studies cannot be translated and is thus not relevant to liposomaldrug delivery.

Microfluidic technologies have been attempted to alleviate some of theshortcomings of conventional bulk-scale liposome production methods(Andar, A.U. et al. (2014) “Microfluidic Preparation of Liposomes toDetermine Particle Size Influence on Cellular Uptake Mechanisms” Pharm.Res. 31:401-13; Hood, R. R. et al. (2014) “Microfluidic-EnabledLiposomes Elucidate Size-Dependent Transdermal Transport,” PLoS One9:e92978). Controlled liposome formation utilizing a microfluidichydrodynamic flow-focusing (MHF) technique substantially decreases sizevariance, with fewer processing steps. Compared to conventionalbulk-scale techniques, traditional MHF methods provide nanoparticleswith enhanced properties including adjustable, narrowly distributeddiameters (Jahn, A. et al. (2004) “Controlled Vesicle Self-Assembly inMicrofluidic Channels with Hydrodynamic Focusing,” J. Am.Chem. Soc.126:2674-2675; Jahn, A. et al. (2007) “Microfluidic DirectedSelf-Assembly of Liposomes of Controlled Size,” Langmuir 23:6289-6293;Jahn, A. et al. (2008) “Preparation of nanoparticles by continuous-flowmicrofluidics,” J. Nanoparticle Res. 10:925-934; Jahn, A. et al. (2010)“Microfluidic mixing and the formation of nanoscale lipid vesicles,” ACSnano 4:2077-2087), and tunable physiochemical properties (Hood, R. etal. (2013) “Microfluidic synthesis of PEGylated and folatereceptor-targeted liposomes,” Pharm. Res. 30:1597-607). However,realtively low throughput has been achieved by traditional microfluidicsystems, which has constrained MHF methods for use in bulk production(e.g., such as for large scale in vivo studies and preclinical trialswhere larger volumes and higher concentrations are required). Thus,sufficient throughput and nanoparticle concentration have not beenachieved utilizing traditional microfluidic techniques.

Accordingly, there is a need for microfluidic methodologies and systemsthat overcome some or all of the above-noted limitations and/ordisadvantages.

SUMMARY OF THE INVENTION

The present invention relates to microfluidic devices, systems andmethods for on-demand formation of agent-loaded liposomal nanoparticles,with clinically useful doses prepared in minutes rather than hours ordays. The disclosed systems and methods provide for sequential liposomeformation, purification, vesicle functionalization for tissue targeting,and active drug loading. In accordance with disclosed embodiments, amicrofluidic system combines liposome formation with in-line samplepurification and remote drug loading for single step, continuous-flowsynthesis of nanoscale vesicles containing high concentrations of stablyloaded drug compounds and/or other therapeutic or diagnostic agents.Using an on-chip microdialysis element, the system enables rapidformation of large transmembrane pH or ion gradients, followed byimmediate introduction of amphipathic drug for real-time remote loadinginto the liposomes. Thus, the liposomes may be loaded efficiently andeffectively in a continuous flow process enabled by steep transmembraneion gradients, which decrease over time. The microfluidic processenables in-line formation of drug-laden liposomes with drug:lipid molarratios of up to 1.3 or more, and a total on-chip residence time ofapproximately 10 minutes or less, representing a significant improvementover conventional bulk-scale methods which require hours to days forcombined liposome synthesis and remote drug loading. The microfluidicplatform may be utilized to support real-time generation of purifiedliposomal drug formulations with high concentrations of drugs andminimal reagent waste for effective liposomal drug preparation at ornear the point of care. Prior methods have failed to provide active drugencapsulation as part of an on-line liposome synthesis process.

The present invention also relates to microfluidic systems, devices andmethods that enable high throughput synthesis of nanoscale liposomesutilizing high aspect ratio hydrodynamic flow-focusing (HAR-MHF)methods. HAR-MHF systems feature high aspect ratio microchannels inwhich one cross-sectional dimension of the focusing channel (e.g.,width) is significantly greater than a second cross-sectional dimension(e.g., height), as compared to conventional microfluidic devices. Theuse of high aspect ratio channels enables high throughput production ofnanoparticles while also preserving size control and low levels ofpolydispersity (comparable to or exceeding that exhibited by traditionalMHF systems). Such microfluidic-generated liposomes may be produced atspeeds which render it feasible for large scale in vivo experiments,preclinical studies, pilot-scale nanoparticle production, etc.

A microfluidic system for continuous flow synthesis and active loadingof liposomes in accordance an embodiment of the present inventioncomprises a substrate having a sample flow channel having a liposomeformation region, a transmembrane gradient formation region, and anagent loading region. In some embodiments, the substrate is comprised ofa thermoplastic material. The liposome formation region comprises aninlet through which a lipid solution flows and inlets through which abuffer solution flows. The lipid solution and the buffer solutioninteract within the sample flow channel and form a population ofliposomes in a sample buffer solution. The transmembrane gradientformation region is configured to establish a liposome transmembranegradient. The agent loading region comprises an inlet through which afirst agent flows and in fluid connection with the sample flow channel.The first agent is mixed with the liposomes received from thetransmembrane gradient formation region, and actively loaded withinintravesicular spaces of the liposomes.

In some embodiments, the transmembrane gradient formation region is amicrodialysis region. In some implementations, the microdialysis regionis a counterflow microdialysis region. The microdialysis region maycomprise a counterflow channel adjacent to the sample flow channel, anda membrane in between or intermediate the sample flow channel and thecounterflow channel. The membrane permits buffer exchange between thesample flow channel and the counterflow channel and establishes atransmembrane ion gradient.

In some embodiments, the formed liposomes have a median diameter ofbetween about 20 nm and about 500 nm, and in some implementationsbetween about 20 nm and about 100 nm. In some embodiments, the liposomeshave a percent polydispersity of less than about 10%, and in someimplementations less than about 5%.

In some embodiments, the liposome formation region also comprises aninlet through which a second agent flows, whereby the second agent ispassively entrapped within or conjugated to the liposomes duringformation.

In some embodiments, the population of liposomes are formed and activelyloaded in the microfluidic system in less than 1 hour. In someimplementations, the transmembrane gradient formation region effectuatesa shift of pH of said sample buffer solution sufficient to enable activeloading of the first agent within the liposomes and in less than about 5minutes. In some implementations, the first agent is selected from thegroup consisting of an anthracycline, an amphotericin, cytarabine, andchlorpromazine. In some implementations, the first agent is anamphipathic peptide or protein. In some embodiments, the liposomesloaded with the first agent exhibit a drug-to-lipid molar ratio ofgreater than 0.5, in some implementations greater than 2.0.

A microfluidic system having a sample flow channel for continuous flowsynthesis and active loading of liposomes according to an embodiment ofthe present invention comprises a liposome formation region configuredto receive a lipid solution and buffer solution, and form a populationof liposomes in a sample buffer solution, and a microdialysis regiondownstream from and in fluid connection with said liposome formationregion and configured to form a liposome transmembrane ion gradient. Insome embodiments, the system further comprises a drug-loading regiondownstream from and in fluid connection with the microdialysis regionand configured to entrap an agent within the liposomes.

In some implementations, the microdialysis region comprises acounterflow channel adjacent to the sample flow channel, and a membranebetween the sample flow channel and the counterflow channel. Themembrane permits buffer exchange between the sample flow channel and thecounterflow channel for enabling removal of free ions from the samplebuffer solution. In some implementations, the membrane prevents selectedparticles or molecules from passing between the sample flow channel andthe counterflow channel.

The present invention also relates to a method for continuous flowsynthesis and active drug loading of liposomes, comprising the steps of:forming a population of liposomes within a buffer solution in amicrofluidic channel; establishing a transmembrane pH gradient betweenan intravesicular space within the liposomes and the buffer solution;and entrapping an agent into the intravesicular space via directed ionexchange to form agent-loaded liposomes. In some implementations, theliposomes are formed and loaded in less than about 1 hour. In someimplementations, the liposomes are formed and loaded in less than about10 minutes.

The present invention is also directed to a microfluidic device forsynthesis of liposomes. The device comprises a sample flow channel, afirst inlet channel in fluid communication with the sample flow channel,and second and third inlet channels in fluid communication with thesample flow channel. The first, second and third inlet channels convergeat a flow focusing region within the sample flow channel, wherein thesample flow channel has an aspect ratio (height:width) exceeding 20:1 atthe flow focusing region. In some implementations, the aspect ratio is50:1 or greater. In some implementations, the aspect ratio is 100:1 orgreater. The first, second and third inlet channels are preferablyvertically oriented relative to each other.

BRIEF DESCRIPTION OF THE DRAWINGS

The file of this patent contains at least one drawing/photographexecuted in color and that copies of this patent with colordrawing(s)/photograph(s) will be provided by the Office upon request andpayment of the necessary fee.

FIG. 1 is a graphical representation of liposome size distribution andresulting volume distribution for a population prepared by aconventional membrane extrusion method. While nearly half of theliposomes are below 100 nm in diameter, these smaller liposomesrepresent less than 0.5% of the total volume of the sample.

FIG. 2 illustrates graphically measured folate content of liposomesformed with both 10% PEG-PE and 2% folate-PEG-PE added to the lipidmixture prior to flow-focusing. Cryo-TEM images of vesicles formed withand without the added folate are shown inset. Efficient liposomePEGylation was confirmed through separate measurements.

FIG. 3 is a schematic representation of an embodiment of a microfluidicdevice (FIG. 3, panel a) providing PEG-lipid and folate-PEG-lipidmicelle injection for rapid introduction of exogenous ligands intoliposomes (hydrophilic drug encapsulation shown). An exploded view ofthe circled portion of FIG. 3, panel a is shown in FIG. 3, panel b.

FIG. 4 illustrates schematic representations (FIG. 4, panel a, and FIG.4, panel b) of a fabricated membrane dialysis chip according to anembodiment enabling rapid liposome purification and buffer exchange forpH adjustment and downstream active drug loading,

FIG. 5 illustrates graphically exemplary results of on-chip membranedialysis for pH adjustment of ˜3 pH units in under 1 min (FIG. 5, panela), enabling rapid active loading of AO as an amphiphilic drug analog ina downstream drug/liposome mixing channel (FIG. 5, panel b). Drug:lipidmolar ratios greater than 2 were achieved, which is significantly higherthan D/L ratios exhibited by conventional bulk processes.

FIG. 6 illustrates schematic representations of a high-throughputnanoliposome drug synthesis chip according to an embodiment of thepresent invention, comprising a radial array (FIG. 6, panel b) of 128flow-focusing elements (FIG. 6, panel a). The synthesis chip includes aspiral element supporting buffer exchange, active drug loading, andpurification downstream of the parallel liposome formation zones. Afluid routing manifold containing flow splitters is aligned to theflow-focusing chip to connect fluid input lines to the array elements,and a single liposome output port delivers purified liposomal drugs to acollection vial (FIG. 6, panel c).

FIG. 7 illustrates schematic representations of a fully-integratedmicrofluidic device in accordance with an embodiment of the presentinvention for remote loading of liposomal therapeutic nanomedicinesin-line with liposome synthesis and buffer exchange via microdialysisfor rapid generation of nearly monodisperse, functionalized liposomeswith tunable diameters containing high concentrations of stably loadedcompounds (FIG. 7, panel a). A cross-sectional view of the microfluidicsystem is shown in FIG. 7, panel b, revealing the differing channelheights supporting each process step.

FIG. 8 illustrates schematic representations of components of aPDMS/cellulose hybrid microfluidic device in accordance with embodimentsof the present invention, with a channel for buffer counterflow (FIG. 8,panel a), patterned nanoporous regenerated cellulose dialysis membrane(FIG. 8, panel b), and sample channel for liposome synthesis, bufferexchange, and remote drug loading (FIG. 8, panel c). A photograph of anexemplary fabricated device is shown in FIG. 8, panel d.

FIG. 9 is a graphical representation of exemplary data for on-chipmicrofluidic buffer exchange via membrane dialysis at variouscounterflow pH and flow velocities. Residence times varied from 40 s to80 s, resulting in a ApH of 1.7 to 3.0. Total flow rates varied fromapproximately 7 μL min⁻¹ to 14 μL min⁻¹ (with average linear velocitiesfrom 0.6 cm s⁻¹ to 0.3 cm s⁻¹, respectively).

FIG. 10 illustrates graphically a numerical simulation of ammoniumsulfate (initial concentration 250 mmol L⁻¹) transport in themicrofluidic device to verify adequate ion removal. FIG. 10, panel adepicts the ammonium sulfate concentration throughout the microdialysissegment of the device (channel length scaled by a factor of 10 for morerapid computation). FIG. 10, panel b depicts concentration profile ofammonium sulfate along the sample channel, RC membrane, and counterflowchannel at the exit of the dialysis region for flow velocities varyingfrom 0.3 cm s⁻¹ to 0.6 cm s⁻¹.

FIG. 11 illustrates graphically the relationships between samplevelocity (FIG. 11, panel a) and initial AO concentration (FIG. 11, panelb) on final encapsulated concentration and loading efficiency.Microfluidic-generated liposomes, 80.8 nm in diameter, were formed in aseparate chip for this experiment.

FIG. 12 depicts graphically volume-weighted size distributions fromremote loading of DOX and AO into liposomes in-line with synthesis andmicrodialysis for buffer exchange in comparison to unloaded liposomes(buffer only) generated by the microfluidic device. A cryoTEM image of aDOX-loaded liposome formed in-line with synthesis is shown inset.

FIG. 13 illustrates schematically a comparison of two MHF devices and anHAR-MHF device (FIG. 13, panel a). Numerical simulations (COMSOL) wereconducted comparing ethanol concentration profiles within MHF andHAR-MHF systems with increasing microchannel aspect ratios, andconsequently increasing ethanol concentration profile uniformities. InHAR-MHF, the focusing axis is perpendicular to that of typical MHF,providing the benefits of an extremely high aspect ratio system withdecreased back pressure and simple fabrication processes for robusthigh-throughput synthesis of liposomes. FIG. 13, panel b is a photographof an actual HAR-MHF device consisting of COC plaques and thin (50 μm)films produced using simple fabrication techniques.

FIG. 14 illustrate graphically size distributions (FIG. 14, panel a,panel b, and panel c) and modal diameters (FIG. 14, panel d, panel e,and panel f) of liposomes produced within an exemplary HAR-MHF device inaccordance with the present invention, with microchannel aspect ratio100:1 (FIG. 14, panel a, and panel d) as compared to liposomes producedthrough traditional MHF using devices with microchannel aspect ratios of5:1 (FIG. 14, panel b, and panel e) and 0.5:1 (FIG. 14, panel c, andpanel f) at various FRRs. HAR-MHF upholds the established advantage ofmicrofluidics to produce narrowly distributed populations of liposomeswith tunable diameters.

FIG. 15 illustrates graphically the polydispersity index (PDI) of eachliposome sample produced within the HAR-MHF device with aspect ratio100:1 compared to liposomes produced through traditional MHF deviceswith aspect ratios of 5:1 and 0.5:1 at various FRRs. Under each flowcondition tested, HAR-MHF produced liposomes with a lower PDI thaneither of the MHF devices.

FIG. 16 are COMSOL simulations illustrating the fluid velocity at themidpoint of the channel width within the HAR-MHF and traditional MHFmicrofluidic devices at FRR 20. As the aspect ratio increases, flowvelocity becomes more uniform across the width of the channel (allprofiles normalized to effective channel height and maximum flowvelocity).

FIG. 17 illustrates graphically modal diameters (FIG. 17, panel a) andPDI (FIG. 17, panel b) of liposomes produced through the HAR-MHF devicewith the initial lipid concentration varying from 5 mM to 80 mM. Anincrease in initial lipid concentration, increase in final liposomeconcentration and production rate, does not largely affect liposomemodal diameter or PDI when increasing from 5 mM up to 40 mM. Higherconcentrations (60 mM and 80 mM) cause an increase in both liposome sizeand polydispersity.

FIG. 18 illustrates graphically modal diameter (FIG. 18, panel a) andPDI (FIG. 18, panel b) of populations of liposomes generated via HAR-MHFat FRR 20 and total linear flow velocities ranging from 0.05 m/s to 0.3m/s (0.75 mL/min to 4.5 mL/min, respectively). For flow velocitieshigher than 0.1 m/s, liposome diameter and PDI are both nearly constant,demonstrating that higher throughput for the HAR-MHF system may beachieved by increasing total flow velocity without causing significantchanges in the resulting liposome populations.

FIG. 19 illustrates graphically a comparison of liposome productionrates across tested microfluidic methods, using a consistent 20 mMstarting lipid concentration. HAR-MHF enables the generation ofliposomes at ˜100 mg/h lipid, two orders of magnitude higher thantraditional MHF, without sacrificing any of the benefits provided bymicrofluidic liposome synthesis. Similarly, HAR-MHF produces liposomesone order of magnitude faster than a capillary-based 3D-MHF method,while providing approximately 10²˜10³ times higher vesicle concentrationthan the capillary system.

DETAILED DESCRIPTION OF THE INVENTION

The present invention relates to systems, devices and methods forsequential liposome formation, purification, and active drug loadingbased on the use of a transmembrane pH or ion concentration gradientestablished using an on-chip microdialysis membrane. The microfluidicsystem provides single-step continuous-flow synthesis of vesiclesencapsulating both lipophilic and amphipathic drugs with minimal reagentwaste. The liposomes may have a selected size, for example havingvesicle diameters of about 20 nm to about 300 nm, with the selectedpopulation size exhibiting extremely low levels of polydispersity. Inaccordance with some disclosed embodiments, systems and methods areprovided that enable high throughput synthesis of nanoscale liposomes.

Disclosed methods and systems enable controllable formation of small anduniform liposomes, thereby allowing for an accurate determination of theimpact of relatively small liposomes (e.g., with vesicles less than 100nm, and as small as 20 nm) for drug delivery applications. Large sizevariability of liposome populations can have a significant negativeimpact on toxicity. As noted above, liposomal drugs prepared bytraditional methods exhibit skew normal size distributions with largevariance, even after multiple processing steps (e.g., membrane extrusionor gel filtration) to improve size homogeneity (see Maguire, L. A. etal. (2003) “Preparation of small unilamellar vesicles (SUV) andbiophysical characterization of their complexes withpoly-1-lysine-condensed plasmid DNA,” Biotechnology and AppliedBiochemistry 37:73-81). For example, size variation data in a liposomepopulation prepared by a conventional 5-step extrusion process is shownin FIG. 1. The left-hand curve presents the distribution of measuredliposome diameters, while the right-hand curve reflects the resultingliposome volume distribution. While 43% of the measured liposomes arebelow 100 nm in diameter, liposomes larger than 100 nm account for morethan 99.5% of the total volume (drug dose). Similarly, although the modeof the size distribution is 95 nm in measured liposomes, the mode of thevolume distribution is 204 nm in total volume. Thus, as demonstrated inFIG. 1, a fundamental problem with liposomal carriers prepared bytraditional bulk synthesis methods is large size variance. Even moderatevariance in liposome size results in a large amount of encapsulated drugbeing introduced by vesicles outside of the desired size range,necessitating the delivery of a larger total drug amount to achieve agiven therapeutic index from vesicles within the optimal range, thusresulting in potential increase in toxicity and degraded safetyprofiles.

Controlled liposome formation utilizing microfluidic hydrodynamicflow-focusing (MHF) substantially decreases size variance (Jahn, A. etal. (2004) “Controlled Vesicle Self-Assembly in Microfluidic Channelswith Hydrodynamic Focusing,” J. Am. Chem. Soc. 126:2674-2675; Jahn, A.et al. (2007) “Microfluidic Directed Self-Assembly of Liposomes ofControlled Size,” Langmuir 23:6289-6293). A much higher percentage ofvesicles within a selected and optimal range may be achieved, therebylowering total drug amount required for the same therapeutic index andsignificantly improving safety profiles. In MHF techniques, liposomesare formed by a diffusively driven process wherein a stream of lipidsolvated in an alcohol is hydrodynamically sheathed between two obliqueaqueous streams within a microfluidic channel. The lipid stream isfocused into a narrow sheet with a thickness varying from a fewmicrometers to several hundred nanometers, depending on the lipid:bufferflow rate ratio. The laminar flow conditions facilitate highlycontrollable diffusive mixing at the two miscible liquid interfaces,diluting the alcohol concentration below the lipid solubility limit andinitiating lipid self-assembly into small unilamellar vesicles. Forexample, a central stream of phospholipids dissolved in 2-propanol (IPA)may be focused by outer streams containing a narrow inner sheath ofgreen carboxyfluorescein dye as a model encapsulant, with an outersheath of aqueous buffer used to define the degree of focusing.

In contrast to other microfluidic techniques based on electroformation(e.g., see Kuribayashi, K. et al. (2006) “Electroformation of giantliposomes in microfluidic channels,” Measurement Science and Technology3212 (2006) or crossflow injection (Wagner, A. et al. (2002) “Thecrossflow injection technique: An improvement of the ethanol injectionmethod,” J. Liposome Res 12:259-270), MHF techniques generate a large,controllable, and spatially-varying solvent gradient, which is believedto be responsible for the formation of highly uniform vesicles withaverage diameters that can be adjusted by simply modifying the flow rateratio. Liposome populations with average liposome diameters ranging fromabout 50 nm to 150 nm have been achieved with relative standarddeviations of about 10-15% utilizing traditional MHF methods (see Jahn,A. et al. (2007) “Microfluidic Directed Self-Assembly of Liposomes ofControlled Size,” Langmuir 23:6289-6293).

Disclosed embodiments of microfluidic flow focusing systems of thepresent invention further extend the selected vesicle size range fromabout 20 nm to about 500 nm, while dramatically reducing polydispersityin the resulting liposome populations to less than about 10%, morepreferably less than about 5%. The disclosed methodologies and systemstherefore allow for precise comparisons of in vitro and in vivoperformance over a wide vesicle size range without the confoundinginfluence of high polydispersity.

Early microfluidic flow-focusing chips were fabricated using acumbersome silicon/glass process. In contrast, embodiments of thepresent invention provide for a robust thermoplastic microfabricationmethod that supports rapid prototyping of different device designs,improves vesicle size range and greatly reduces polydispersity. A fullyintegrated liposomal drug synthesis process is provided, which includesvesicle formation, lipophilic and amphipathic drug encapsulation,vesicle functionalization for molecular targeting, vesicleconcentration, and liposomal drug purification. The thermoplasticmicrofabrication method supports low-cost scale-up of the technology forproducing clinically-acceptable volumes of targeted liposomal drugs in asingle continuous-flow process. Thus, a fully integratedpharmacy-on-a-chip platform is achieved, enabling production of a newgeneration of size-optimized, targeted and multi-agent liposomal drugs.

The microfluidic process of the present invention may be implemented forfacile production of custom liposomal drugs for preclinical or clinicalresearch applications, scaled up as part of a larger scale productionpipeline, or implemented for point-of-care applications. Liposomesdispersed in aqueous solution are subject to physical and chemicalinstabilities during long-term storage, including lipid hydrolysis andoxidation, liposome aggregation, and liposome fusion (see Chen, C. etal. (2010) “An overview of liposome lyophilization and its futurepotential,” J. Controlled Release 142:299-311). Loss of encapsulateddrugs due to membrane leakage (e.g., such as during long-term storage)reduces the therapeutic efficiency of liposomal drugs over time. Forexample, storage of siRNA-loaded liposomes at 4° C. for one month hasbeen observed to result in significant encapsulant leakage and an 80%decrease in gene silencing efficiency (Chang, H. I. & Yeh, M. K. (2012)“Clinical development or liposome-based drugs: formulation,characterization, and therapeutic efficacy,” International J. ofNanomedicine 7:49-60). As such, conventional liposomal drug suspensionsare typically stated as having shelf lives ranging from 12 to 20 months,with lipid species selected for enhanced stability rather than for theirtherapeutic efficacy. Even in the case of Doxil®, which is subject tominimal leakage since the majority of the encapsulated doxorubicin isstored in a crystallized state, up to 10% of the encapsulated drug canleak from the vesicles during storage over its 20 month shelf life.

Furthermore, the size distributions of stored liposomal drugs changeover time due to vesicle fusion. While lyophilization of liposomes canavoid this latter concern and further extend storage times, the freezedrying process itself results in significant drug leakage and vesiclefusion (Chen, C. et al. (2010) “An overview of liposome lyophilizationand its future potential,” J. Controlled Release 142:299-311). Themicrofluidic process of the present invention offers an innovativeapproach to preparing on-demand nanoliposome drugs for almost immediateuse in experimental or point-of-care settings. As such, lipidcompositions and bioactive agents tailored for optimal performance maybe utilized without regard for long-term stability. Point-of-careliposomal drugs may be prepared in accordance with the disclosed systemsand methodologies, with targeting ligands customized to a patient'sspecific disease profile.

The microfluidic liposome formation process of the present inventionprovides for dynamic control over vesicle size. The microfluidicflow-focusing technique is optimized to produce smaller and more uniformliposome populations compared to those achievable using prior bulk ortraditional microfluidic methods. Using a high aspect ratio microchannelfabrication method, the formation of nearly-monodisperse liposomepopulations with a selected size within a wider size range and withlower size variance is achieved.

In prior preparation methods utilizing a bulk alcohol injection process,a stream of lipids in alcohol is injected into a vortexed aqueous bufferunder high Reynolds number flows. However, the bulk alcohol injectionprocess does not allow for controlled mixing conditions or solventgradients, and is sensitive to perturbations in the mixing conditionswhich result in large shifts in mean liposome diameter and sizevariance. In contrast, the techniques disclosed herein afford exquisitecontrol over the local flow conditions, and enable the formation ofsmall liposomes with tightly controlled distributions.

Based on the disclosed studies of the present invention, liposome sizeis believed to be inversely related to the aqueous:alcohol flow rateratio, and remains independent of the total flow rate. Theinterpretation of this relationship is that at higher flow rate ratiosthe focused lipid stream is narrowed, resulting in a more rapid decreasein alcohol concentration within the focused stream. As a result, thetime scale between initial formation of single lipid layers, or lipid“leaflets,” and the collapse of leaflets into closed and fully assembledliposomes is reduced, resulting in reduced leaflet growth and thussmaller vesicles. Maintaining uniform focusing of the lipid stream ateach channel cross-section thus achieves uniform liposome populations,since variations in the stream width lead to variations in diffusiontime scales and thus wider liposome size distributions.

Early flow-focusing chips fabricated in etched silicon substrates hadchannel aspect ratios (height:width) that were substantially uniform(i.e., ˜1). Due to the parabolic flow profile associated withpressure-driven microchannel flow in such early chips, with no-slipconditions imposed at the channel walls, the lipid stream at the upperand lower surfaces of the focusing channel is not focused by the aqueoussheath flows, and thus contributes to unwanted variance in the sizedistribution. Furthermore, differences in focused stream width increasewith the flow rate ratio, leading to higher variance for smallerliposomes and placing a limit on the minimum vesicle size that can begenerated using such early techniques. For example, by using a siliconeelastomer soft lithography process supporting a 4:1 channel aspectratio, a wider range of liposome sizes ranging from about 40 nm to about277 nm was achieved with variance below 8% over the full size range.

In accordance with embodiments of the present invention, vesicle sizecontrol is further improved by implementing a dry film photoresistthermoplastic microfabrication process capable of providingsignificantly higher aspect ratio channels than those previouslyattempted. Thermoplastic cyclic olefin copolymer (COC) is a preferredsubstrate material for the flow-focusing chip due to its low materialcost, compatibility with patterning by rapid replication methods such asroll-to-roll hot embossing, excellent solvent compatibility, andcapabilities for heterogeneous material integration. While thermoplasticmicrochannels are typically limited to low aspect ratios, thislimitation has been overcome through the use of a multi-layer dry filmphotoresist patterning process.

Dry film resists (DFRs) are thick sheets of photopatternable epoxy-basedresist sometimes employed for circuit board patterning, but which havenot previously been utilized in nanoparticle synthesis. Unlikephotopatterned molds fabricated from thick spin-coated photoresists(e.g., SU-8), the continuous DFR sheets are applied by lamination andthus multiple layers can be applied even after patterning the underlyingfilms. Furthermore, unlike silicon molds, the tough DFR molds do notreadily fracture during the embossing process. This combination offeatures makes DFR molds preferred for replica molding of high aspectratio thermoplastic chips. For example, individual 30 μm thick DFRlayers can be patterned with about 10 μm to about 15 μm lateral featuresfor an aspect ratio of at least 2:1. The photolithography process may befurther optimized for sequentially processing 5 or more DFR layers foran overall aspect ratio of 10:1 or greater. Using this process, a widerrange of liposome sizes may be achieved, for example ranging from about20 nm to about 500 nm, more preferably from about 20 nm to about 300 nm,more preferably from about 20 nm to about 100 nm, and with low sizevariation, preferably less than about 10%, more preferably less thanabout 5%. Thus, the disclosed methodologies represent anorder-of-magnitude reduction in polydispersity over prior bulk membraneextrusion processes (e.g., see Berger, N. et al. 2001) “Filter extrusionof liposomes using different devices: comparison of liposome size,encapsulation efficiency, and process characteristics,” InternationalJournal of Pharmaceutics 223:55-68). Moreover, the relationships betweenliposome size distributions and system-level parameters that areaccessible to the chip designer may be optimized as desired, includingchannel width, channel aspect ratio, and flow-focusing intersectiongeometry.

The flow-focusing process has also been extended to a wider range oflipids and lipid formation conditions than previously attempted. Inprevious methods, liposomes were formed using the neutral lipiddimyristoyl phosphatidylcholine (DMPC), combined with cholesterol anddihexadecyl phosphate (DCP) in a molar ratio of 5:4:1 (Kuribayashi, K.et al. (2006) “Electroformation of giant liposomes in microfluidicchannels,” Measurement Science and Technology 17:3121; Jahn, A. et al.(2010) “Microfluidic mixing and the formation of nanoscale lipidvesicles,” ACS nano 4:2077-2087; Hood, R. et al. (2013) “MicrojluidicSynthesis of PEG- and Folate-Conjugated Liposomes for One-step Formationof Targeted Stealth Nanocarriers,” Pharm. Res. 30:1597-1607). In thepresent invention, the flow-focusing process may be utilized with otherneutral lipids with different fatty acid chain lengths (DOPC and DLPC),as well as an anionic lipid (DOPC).

In accordance with embodiments of the present invention, a microfluidicsystem is provided that enables rapid and efficient active or remoteloading of an agent(s) or drug(s) into nanoscale liposomes, combiningliposome synthesis and remote drug loading in a continuous integratedprocess. Conventional bulk methods for remote drug loading require aseries of discrete manual operations using large fluid volumes. Inaddition, prior microfluidic methods, although providing some advantagesover conventional bulk preparation methods, have been limited to passivedrug encapsulation and require off-chip sample purification to removeresidual solvents or non-encapsulated reagents. During passive loading,the desired drug is added to the lipid mixture prior to vesicleformation, whereby the drug becomes sequestered within the liposomesduring the self-assembly process. However, passive encapsulation isinefficient, with less than 10% encapsulation efficiencies typicallyachieved for hydrophilic compounds, resulting in significant waste ofvaluable drug. In addition, the maximum attainable D/L ratio duringpassive loading is limited by drug solubility (Cullis, P. R. et al.(1989) Adv. Drug Delivery Rev. 3:267-282), constraining the total amountof drug that can be encapsulated. Further, in the case of amphipathicdrugs that possess both hydrophilic and hydrophobic regions, drugencapsulated by prior methods ultimately migrates out of the vesicles,resulting in varying concentration levels over time (Abraham, S. A. etal. (2005) Methods Enzymol., 391:71-97).

In contrast, liposome formation and drug-loading techniques inaccordance with disclosed embodiments provide for a counterflowmicrodialysis element, which enables steep transmembrane pH or iongradients to be formed immediately prior to active or remote drugloading of the formed liposomes. In some embodiments, a microfluidicdevice includes a sample flow channel including a continuous flow paththat proceeds from a liposome formation region to a microdialysisregion, and then to a drug-loading region (e.g., see FIG. 7, panel a).The liposome formation region may include an inlet through which a lipidsolution is injected, as well as inlets through which buffer solutionsare injected, followed by and in fluid communication with a liposomestabilization channel. In some implementations, a first agent or drugmay be introduced in the liposome formation region for passiveencapsulation during the assembly process (e.g., see FIG. 3).

The formed liposomes in sample buffer solution then proceed through amembrane dialysis or buffer exchange region. A counterflow channel isprovided in the membrane dialysis region, through which a dialysisbuffer solution flows in a direction opposite to or against thedirection of flow of the sample buffer solution (e.g., see FIG. 7, panelb). The pH of the sample buffer solution is less than the pH of thedialysis or counterflow buffer solution. A membrane is disposed betweenthe sample flow channel and the counterflow channel, permitting bufferexchange between the sample flow channel and the counterflow channel andthus establishing a liposomal transmembrane pH gradient.

Immediately after formation of the transmembrane gradient, the liposomesflow into the drug loading region where an agent or drug is introducedand encapsulated (see FIG. 7, panel a). Drug encapsulation via active orremote loading takes advantage of the transmembrane chemical gradientsto entrap the agent from the surrounding environment (i.e., theextravesicular spaces) into the already formed liposomes (i.e. withinthe intravesicular spaces). During the remote loading process, the agentor drug (e.g., amphipathic drug) diffuses through the bilayer lipidmembrane into the intravesicular space within the liposome. Once insidethe liposome, a chemical modification of the drug occurs, preventingmembrane repermeation and thereby resulting in the accumulation of drugwithin the liposomes.

Some embodiments of microfluidic devices or chips provide for active orremote drug loading using a pH gradient to encapsulate the drug (e.g.,amphipathic weak bases). In this approach, liposomes are initiallyformed in an acidic environment. After vesicle self-assembly, theinterior of the liposome remains acidic while the extravesicular pHlevel is adjusted to physiological conditions. Incubation with unchargeddrug allows molecules to diffuse into the liposomal intravesicularcavity, where the drug molecules then become protonated. The positivelycharged drug molecules can no longer traverse the bilayer membrane andare thus effectively trapped inside the liposomes.

Other chip embodiments provide for remote loading using a transmembraneion gradient, for example for loading of amphipathic weak bases and/oracids. In this approach, liposomes are formed with a high concentrationof a suitable ionic species selected to act as a counterion to theamphipathic drug. As the drug crosses the liposome membrane, it israpidly formed into an insoluble salt through ionization, resulting inthe formation of an insoluble salt which cannot diffuse back into theextravesicular environment, resulting in exceptionally high loadinglevels and improved liposome stability during storage and circulation.

Other chip embodiments provide for a combination of passive and activeloading in a continuous flow process. An integrated microfluidic chip isprovided for vesicle formation, PEGylation, targeting moleculeattachment, and multi-agent encapsulation of amphipathic and lipophilicdrugs, with a first agent entrapped during liposome formation viapassive loading, and additional and/or a second agent entrapped withinthe formed liposomes via active loading. For passive loading, tertiarychannels may be provided through which solvated compounds are injectedfor encapsulation during liposome formation. The differential solubilityof lipophilic drugs directs diffusion toward the center of the streamwhere vesicle formation occurs.

Any of the disclosed systems and methods may be scaled up forlarge-volume and/or on-demand production, such as for point-of-care andclinical dosing levels. In accordance with embodiments of the presentinvention, a high aspect ratio microfluidic hydrodynamic flow-focusing(HAR-MHF) system is provided featuring high aspect ratio microchannelscomposed of extremely wide channel widths and relatively narrow channelheights, enabling high throughput production of nanoparticles whilepreserving size control and low polydispersity. In some implementations,liposomal synthesis techniques are carried out utilizing devices havingmicrochannels with unique flow focusing axis orientation. Due to theorientation of the flow focusing axis, HAR-MHF combines the advantagesof traditional MHF with the benefits of also providing extremely highaspect ratios for exceedingly uniform flow profiles and amplifiedproduction rates. Governed by the same physics as MHF for controlledliposome self-assembly, HAR-MHF reduces the necessary amount ofparallelization, thus further increasing sample homogeneity and ease ofoperation. Nearly monodisperse liposomes of tunable size may begenerated at unprecedented rates, thereby providing microfluidicliposome production suitable for large scale in vivo and preclinicalapplications.

Embodiments of the present invention also provide for functionalizationof liposomes, such as via PEGylation. Conventional chemotherapeuticsexhibit poor specificity in reaching tumor tissues. The location andextent of metastatic tumors limit their accessibility, requiringsystemic administration which increases drug toxicity. Targeted deliveryof nanoparticle-based drugs may be utilized using a variety oftumor-selective ligands including antibodies, aptamers, and peptidesconjugated to the nanoparticle surfaces (e.g., see Gu, F. X. et al.(2007) “Targeted nanoparticles for cancer therapy,” Nano Today 2:14-21).While these targeting agents can offer high specificity, they arecomplex and expensive to manufacture using conventional methods, withhydrodynamic radii that can be of the same order of magnitude as thenanoparticles themselves, thus limiting the number of molecules that canbe attached for targeted delivery and affecting in vivo behavior.

An alternative small molecule ligand for tumor targeting is folate(Yoshida, T. et al. (2006) “Induction of cancer cell-specific apoptosisby folate-labeled cationic liposomes,” J. Controlled Release111:325-332). Folate receptors are highly-overexpressed in a wide rangeof cancers, including tumors of the ovary, brain, kidney, lung andbreast (Garin-Chesa, P. et al. (1993) “Trophoblast and ovarian cancerantigen LK26, Sensitivity and specificity in immunopathology andmolecular identification as a folate-binding protein,” Am. J. Pathol.142:557-567), with up-regulation levels that tend to correlate withtumor stage (Toffoli. G. et al. (1997) “Overexpression of folate bindingprotein in ovarian cancers,” International J. Cancer 74:193-198). Folatebinds to these receptors with exceptionally high affinity (Lee. R. J. &Low, P. S. (1994) “Delivery of Liposomes into cultured KB cells viafolate receptor-mediated endocytosis,” J. Biol. Chem. 269:3198-3204),making it an excellent candidate for targeted delivery. Folate may bereadily conjugated with liposomes using polyethyleneglycol (PEG) linkerswithout affecting its activity (Lee. R. J. & Low, P. S. (1994) “Deliveryof liposomes into cultured KB cells via folate receptor-mediatedendocytosis,” J. Biol. Chem. 269:3198-3204).

Conventional methods for bulk preparation of folate-functionalizedliposomes is normally performed by adding a defined concentration offolate-polyethylene glycol (PEG)-lipid conjugates to an initial lipidsolution prior to vortexing and filtration to form the liposomes. Inaccordance with methods of the present invention, phosphoethanolamine(PE)-containing lipids conjugated with PEGs000 are added to a DMPC lipidmixture. Little variation in liposome diameter was observed when PEG-PEwas added to the lipid solution, despite the large size of the PEGs000molecules. Next, 2 mol% folate-PEG-DSPE(1,2-distearoyl-sn-glycero-3-phosphoethanolamine-N-(folate(polyethyleneglycol)₂₀₀₀)) is added to the lipid solution. FIG. 2 reveals successfulincorporation of folate into the resulting liposomes, as determined byUV/vis absorption measurements from the purified liposomes. Thus, bothPEGylated and folate-functionalized liposomes may be successfullyproduced in the microfluidic platform.

While pre-insertion methods offer a simple procedure for producingfolate receptor-targeted liposomal drugs, a disadvantage of thisapproach is that the resulting liposomes present PEG and folate on bothinner and outer lipid leaflets. This is a particular concern whenforming smaller liposomes, since the large PEG molecules on the innerleaflet limit the internal volume available for drug encapsulation. Toavoid this concern, a post-insertion route to folate functionalizationmay be utilized. As shown in FIG. 3, PEG-lipid and folate-PEG-lipidconjugates are introduced through parallel side channels atconcentrations above their critical micelle count, with hydrophilic orlipophilic drug injection channels positioned between the main lipidchannel and micelle injection channel. The small drug molecules rapidlydiffuse into the liposome formation region and become encapsulated,while the larger micelles with a longer diffusion length scale reach thepre-formed liposomes downstream of the initial lipid mixing region. Insome implementations, a thermoelectric heating element is integratedinto the chip to provide local temperature control, and raise theliposomes above the lipid phase transition temperature to encourage theinsertion of PEG and folate-PEG into the outer monolayer of thevesicles.

To increase the residence time of interacting micelles and liposomes, along and wide serpentine outlet channel may be provided. However, due tothe short diffusion length scales on the order of several hundrednanometers, equilibrium surface ligand concentrations are rapidlyreached within the flow system. Thus, the kinetics of PEG and folate-PEGincorporation may be further optimized by varying channel geometries,flow conditions, and micelle concentrations.

Membrane Dialysis and Active Amphipathic Drug Loading

Following liposome synthesis, the nanoparticles may be purified toremove free lipophilic drug not encapsulated into the vesicles. Forconventional bulk liposome preparation, size exclusion chromatography(gel filtration) is the preferred method for vesicle purification (seeHolzer, M. et al. (2009) “Preparative size exclusion chromatographycombined with detergent removal as a versatile tool to prepareunilamellar and spherical liposomes of highly uniform sizedistribution,” J. Chromatography A 1216:5838-5848). However,implementation of this approach in a continuous flow-through system isnot feasible since separate loading and elution buffers must besequentially applied to the chromatography column. Furthermore, the useof a packed bed of gel chromatography media introduces high hydrodynamicresistance that is incompatible with the flow rates required forliposome formation. Gel filtration also results in vesicle dilution, asignificant disadvantage for pharmacy-on-a-chip applications.

In accordance with embodiments of the present invention, the continuousflow system of the present invention utilizes an on-chip membranedialysis element. The membrane dialysis technique achieves efficient andrapid liposome purification and buffer exchange in a compactflow-through format, and without any vesicle dilution. In someimplementations, a nanoporous cellulose membrane is permanently sealedby solvent bonding between two polymer substrates containingmicrochannels. The ion-permeable membrane allows buffer exchange betweenthe liposome sample flow injected through the top channel and acounterflow buffer injected through the bottom channel in opposition tothe liposome sample flow, thereby maximizing average solute gradients.

In addition to providing a simple flow-through method for liposomepurification that is directly compatible with the flow-focusingtechnique, the on-chip membrane dialysis element enables highlyefficient and continuous-flow active loading of an agent or drug intothe preformed liposomes. Any desired agent or drug suitable forliposomal encapsulation may be utilized, including but not limited toanthracyclines (doxorubicin, daunorubicin, aciacinomycin, etc,),amphotericin, cytarabine, chlorpromazine, and/or amphipathic peptide orprotein drugs (with both hydrophobic and hydrophilic groups, which aresoluble in water but also traverse the lipid bilayer into the liposomecore).

As a result of the buffer exchange, a transmembrane pH gradient isestablished, so that amphipathic weak acids and bases are activelyloaded into preformed vesicles via directed ion exchange. The activeloading enables high drug concentrations to be entrapped within thevesicles with excellent long-term stability. In comparison toconventional active loading methods that require hours or days forincubation at elevated temperatures, utilization of the membranedialysis element in the disclosed systems quickly shift the pH of theliposome buffer and thereafter actively load the liposomes in minutes(e.g., less than 30 minutes, more preferably less than 10 minutes, morepreferably less than 5 minutes, more preferably less than 3 minutes).For example, in some implementations, the pH of the liposome buffer isincreased by 3 pH units, such as from an initial level of pH ˜6 to ashifted level of pH ˜9. The high transmembrane pH gradient allows aselected agent (e.g., such as doxorubicin, an amphipathic weak base, pKa8.3) to be rapidly driven into the aqueous liposome core.

An exemplary configuration of a membrane dialysis element of amicrofluidic device or chip is illustrated in FIG. 4. Upstream from andin fluid communication with the membrane dialysis element or region is aliposome formation and flow focusing region configured for liposomeformation. Downstream from and in fluid communication with the membranedialysis region is a mixing or loading region configured for receivingan agent (e.g., amphipathic drug) after pH adjustment in the dialysisregion. The dialysis region includes an upper or liposome sample flowchannel, and a lower or counterflow channel (FIG. 4, panel b).

In one implementation, liposomes were formed in a buffer solution at pH4.6. A counter-flow buffer at pH 9.6 was applied to the counterflowchannel. Measurements of pH at the liposome sample flow channel outletas a function of residence time within the dialysis zone revealed shiftsof about 3 pH units in less than 1.5 min (FIG. 5, panel a), sufficientto enable active loading of a desired agent (e.g., acridine orange-HCL(AO; pKa 10.4) as an amphipathic drug analog) immediately after bufferexchange in the dialysis region. Optical absorption measurementsrevealed entrapment into the microfluidic-synthesized liposomes atdrug:lipid molar ratios ranging from 1.6 to 2.3 (FIG. 5, panel b), farexceeding typical D/L ratios of below 0.4 exhibited by conventional bulkactive loading methods which require significantly longer processingtimes (e.g., on the order of several hours or longer).

Because transmembrane pH gradients are unstable, loading efficiency issubstantially improved by the disclosed methods by reducing the timebetween buffer exchange and introduction of the agent or drug (e.g.,amphipathic compound). The inherently decreased diffusion lengths inmicrofluidics enable rapid microdialysis and significantly reducedincubation times for drug loading. To further enhance loadingefficiency, a simple passive herringbone mixing zone (e.g., see Du, Y.et al. (2010) “A simplified design of the staggered herringbonemicromixer for practical applications,” Biomicrofluidics 4:1-13) may beadded downstream of the membrane dialysis region to promote effectivedrug/vesicle interactions. The high loading efficiency of the disclosedtechniques obviates the need for post-entrapment purification (e.g., seeFenske, D. B. & Cullis, P. R. (2010) Liposome Technology, Vol. IIEntrapment of Drugs and Other Materials into Liposomes), achievingexceptionally high efficiency as compared to conventional bulkprocessing methods. However, if further purification is desired, asecondary membrane dialysis region may be added to the system followingthe initial loading region.

Liposome Concentration Enhancement

The resulting drug-loaded liposomes prepared in accordance withdisclosed embodiments exhibit clinically-useful drug concentrationlevels. For doxorubicin, a concentration of 2 mg/mL was targeted,corresponding to a requirement of between 10¹⁴˜10¹⁵ liposomes/mLassuming monodisperse 30-100 nm vesicles with crystallized doxorubicinfilling 50% of the aqueous core.

The flow-focusing process operates at typical bulk flow rates of up to˜100 μL/min, with average flow velocities approaching 10 cm/s. While thecorresponding Reynolds numbers tend to be small, on the order of Re˜10⁻², the flows are characterized by high Peclet numbers which are welloutside the Taylor dispersion regime for even the smallest liposomes ofinterest. Thus, longitudinal convection dominates vesicle transport, andthe stream of liposomes remains tightly focused at the center of thechannel over long length scales relative to the lateral channeldimensions. This characteristic of the system is leveraged to improveliposome concentration by reducing the width of the outer sheath flow atthe liposome formation site. This enables a proportional reduction inthe total volumetric buffer flow rate without impacting liposomeformation rate, thereby leading to an equivalent increase in vesicleconcentration.

In preliminary studies, flow-focusing elements with 100 μm wide sheathchannels yielded vesicle concentrations on the order of 10¹³liposomes/mL, below the target concentration. However, reducing thechannel width to 10 μm provided a ten-fold increase in downstreamliposome concentration to 10¹⁴ liposomes/mL. In preliminary studies,experiments were performed using lipid concentrations more than an orderof magnitude below critical micelle concentration limits, thusindicating that significantly higher concentrations of solvated lipidsmay be used in the flow-focusing process. In further studies, the finalliposome concentration was increased by another order of magnitude to10¹⁵ liposomes/mL by modifying the flow-focusing geometry while alsousing higher initial lipid concentrations.

Process Scale-Up

Whether targeted for point-of-care application or industrial drugproduction, the disclosed pharmacy-on-a-chip systems and methodsdemonstrate the ability to generate sufficient quantities of drug-ladennanoliposomes for clinical use. For the case of liposomal doxorubicinm,a single 50 mg dose requires ˜10¹⁵ liposomes. Using prior flow-focusinggeometry and flow conditions, typical liposome production rates are onthe order of 10¹² liposomes/min. Thus, such conventional methods wouldrequire 10³ minutes (about 16 hours) to produce sufficient drug-loadedliposomes for a single drug dose.

In order to generate sufficient quantities of liposomes for clinicalapplications, throughput may be increased by over two orders ofmagnitude via the parallel operation of array elements. An exemplaryhigh-throughput nanoliposome drug synthesis chip in accordance with anembodiment of the present invention is shown in FIG. 6. The microfluidicchip contains a radial array of 128 individual flow-focusing elements,with inlets connected to a set of single fluid ports through flowsplitter networks. Two of the flow splitters, feeding the lipid andlipophilic drug inlets, are fabricated in an upper routing manifold andconnected to the flow-focusing array through vertical feedthroughs,while a third flow splitter feeding the buffer inlets is fabricated inthe same substrate as the flow-focusing array. In some implementations,all flow splitters and flow-focusing elements are integrated in a singlesubstrate, using a thermoplastic fabrication process that employs a dryfilm photoresist molding template that can create up to 8 individualfluidic layers with photolithographically-patterned vertical passagesfor multilayer interconnection. Flow-focusing channel dimensions may beselected to support vesicle concentrations on the order of 10¹⁴-10¹⁵liposomes/mL. The outlets of the parallel flow-focusing elements arecoupled with a downstream spiral channel supporting sequential bufferexchange to introduce ammonium salt surrounding the vesicles, followedby active drug loading driven by the resulting transmembrane pHgradient, and (if desired) a secondary dialysis region to remove anyremaining free amphipathic drug (or other agent) and yieldultra-purified drug-laden liposomes suspended in PBS buffer suitable forin vivo use.

Parallel operation of the array elements allows the time scale forsingle dose preparation to be reduced to less than 1 hour, preferablyless than 30 minutes, more preferably less than 10 minutes, includingvesicle formation, membrane functionalization, lipophilic drug loading,purification, buffer exchange, and active loading of amphipathic drug.Throughput may be further increased to industrial-scale preparationlevels by operating multiple synthesis chips in tandem.

Having described features and embodiments of the present invention, thesame will be further understood through reference to the followingadditional examples and discussion, which are provided by way of furtherillustration and are not intended to be limiting of the presentinvention.

Device Fabrication

Microchannels were fabricated in polydimethylsiloxane (PDMS) substratesby soft lithography techniques using dry film photoresist molds(Stephan, K. et al. (2007) “Fast prototyping using a dry filmphotoresist: microfabrication of soft-lithography masters formicrofluidic structures,” J. Micromech. Microeng., 17:N69-N74) producedin a multilayer lamination process. Dry film photoresist (Riston MM115i,DuPont, Research Triangle Park, N.C.) was laminated onto a clean glassslide at 110° C. using a feed rate of 0.02 m s⁻¹ and placed on a hotplate at 110° C. for 20 min to promote adhesion. The substrate waspatterned by contact photolithography using a UV flood exposureinstrument (PRX-1000; Tamarack Scientific Co., Corona, Calif.) at a doseof 72 mJ cm⁻² for a single layer of photoresist. Multiple photoresistlayers were processed sequentially using this approach, with a 1.2×increase in UV dose per layer. Each layer of the dry film photoresist isapproximately 37 μm, as measured by stylus profilometry. Following UVexposure, the multilayer substrate was developed using a 1 wt % sodiumcarbonate solution. The resulting molds feature three regions with threedifferent channel heights (FIG. 7, panel a and panel b). Specifically,the flow focusing and liposome stabilization region was 30 μm wide and112 μm deep, the buffer exchange region was 1.2 mm wide and 37 μm deep,and the drug loading and mixing region was 30 μm wide and 77 μm deep.Buffer counterflow channels were 1.2 mm wide and 37 μm deep.

The dry film photoresist molds were next used to create microchannels inPDMS. Two separate molds were used to form an upper substrate containingflow focusing, buffer exchange, and drug loading channels, and a lowersubstrate containing a buffer counterflow channel to assist inmicrodialysis. The molds were placed in plastic petri dishes and a 10:1(w:w) mixture of pre-polymer PDMS elastomer and curing agent (Sylgard184, Dow Corning Corp. Midland, Mich.) was poured on top. Vacuum wasapplied to remove air bubbles, and the petri dish was placed in aconvection oven at 80° C. for 4 h to ensure complete curing of the PDMS.The PDMS substrates were removed from the molds and sectioned using afresh scalpel. Holes for inlet and outlet interfacing were made using amicrobore biopsy punch (Harris Uni-Core, Ted Pella, Inc., Redding,Calif.). All PDMS surfaces were cleaned with isopropanol and DI water.

To form the microdialysis elements, (12 to 14) kDa molecular weight (Mw)cutoff regenerated cellulose (RC) membranes (Spectra/Por 4, SpectrumLaboratories Inc., Rancho Dominguez, CA) were placed between the upperand lower PDMS substrates. The membranes were selected to ensure thatthe nominal pore size (<4 nm) was below the minimum liposome size butlarge enough to allow all individual chemical species and buffer saltsto transport efficiently through the membrane. The RC membranes were cutinto patterns similar to the microchannel geometries using an automatedcraft cutter (Cameo Digital Craft Cutting Tool, Silhouette America,Inc., Orem, Utah), allowing space between adjacent channels for directPDMS-PDMS contact in these regions. The patterned membranes wereflattened using a hydraulic hot press (Carver, Wabash, Ind.) at 0.7 MPafor 10 min at 110° C. prior to chip integration. To enhance sealingbetween the PDMS substrates, a 10:1 (w:w) mixture of pre-polymer PDMSelastomer and curing agent was poured over a glass slide and spin coatedat 3500 rpm for 60 s. The bottom piece of PDMS containing thecounterflow channels was stamped onto the thin layer of PDMS, whichserved as a sealing agent for the microchannels. The patterned RCmembrane was aligned with the microchannels on the top piece of PDMScontaining the flow focusing, dialysis, and drug loading regions. Thetwo substrates were aligned and pressed together by hand, then placed ina convection oven at 80° C. overnight to cure the intermediate PDMSbonding layer. A schematic of an exploded view of the device componentsis shown in FIG. 8, panel a, panel b, and panel c, and a photograph ofthe actual device is shown in FIG. 8, panel d.

Lipid Film and Buffer Preparation

Dimyristoylphosphatidylcholine (DMPC), cholesterol, anddipalmitoylphosphatidylethanolamine-PEG 2000 (PEG2000-PE) (Avanti PolarLipids Inc., Alabaster, Ala.) were combined in chloroform (MallinckrodtBaker Inc., Phillipsburg, N.J.) at a molar ratio of 55:35:10. The lipidmixture was prepared in a glass scintillation vial then stored in avacuum desiccator for at least 24 h for complete solvent removal. Thedesiccated lipid mixture was re-dissolved in anhydrous ethanol (SigmaAldrich, St. Louis, Mo.) for a total lipid concentration of either 40mmol L⁻¹ or 20 mmol L⁻¹, as noted.

Ammonium sulfate (250 mmol L⁻¹, adjusted to pH 4.6) and isosmotic4-(2-hydroxyethyl)-1-piperazineethanesulfonic acid (HEPES) (10 mmol L−1with 140 mmol L−1 sodium chloride, adjusted to pH 7.6) were prepared formicrodialysis and remote loading experiments (both from Sigma-Aldrich).In some cases, trisodium 8-hydroxypyrene-1,3,6-trisulfonate (pyranine)(Invitrogen, Carlsbad, Calif.) was added to the buffers for pHmeasurements (1 μmol L−1). Acridine Orange hydrochloride (AO) at aninitial concentration of 10 mg mL⁻¹ (Sigma-Aldrich) was further dilutedin deionized water as noted and used for remote loading experiments.Doxorubicin hydrochloride (DOX) (Sigma-Aldrich) was diluted to 1.4 mgmL⁻¹ for in-line synthesis and remote loading experiments. All solventsand buffers were passed through 0.22 μm filters (Millipore Corp., NewBedford, Mass.) before being introduced to the microfluidic device.

Numerical Simulation of Ion Exchange Via Microdialysis

Exchange of ammonium sulfate ions during microdialysis was investigatedvia numerical simulations with a two-dimensional model using COMSOLMultiphysics 4.1 (COMSOL, Inc., Burlington, Mass.). The Transport ofDiluted Species (chds) physics interface was applied for the simulationto depict the concentration profiles of the ammonium sulfate salt withinthe microchannels. Microchannel dimensions from the actual devicesfabricated and RC membranes were used to build the model, as well as theknown value of the diffusion coefficient of ammonium sulfate in water atroom temperature (D=8.0×10⁻⁶ cm² s⁻¹).

Buffer Exchange and Remote Drug Loading

Microfluidic devices comprising only the 27 cm long buffer exchange zone(sample channel and buffer counterflow channel) were first used tocharacterize performance of the microdialysis element for rapid ionexchange and remote drug loading. To evaluate buffer exchange, ammoniumsulfate (pH 4.6) was injected into the sample inlet, and isosmotic HEPES(pH 7.6 or pH 9.6) was injected into the buffer counterflow inlet withthe resulting sample collected for analysis. Sample and counterflow flowvelocities were kept equal to one another, and varied from 0.3 cm s⁻¹ to0.6 cm s⁻¹ (approximately 7 μL min⁻¹ to 14 μL min⁻¹, respectively).Pyranine was used as a pH-sensitive molecular probe to determine the pHof the sample and counterflow buffer eluents. Fluorescence intensitymaxima of pyranine at 400 nm and 450 nm is strongly dependent onhydrogen ion concentration, and thus measuring the ratio of the 510 nmfluorescence signal at these excitation wavelengths allows solution pHto be determined (Kano, K. & Fendler, J. H. (1978) “Pyranine as asensitive pH probe for liposome interiors and surfaces. pH gradientsacross phospholipid vesicles” Biochim. Biophys. Acta, 509(2):289-299).Off-chip samples as well as standard curves for calibration over therange from pH 3 to pH 12 were measured using a SpectraMax plate reader(Molecular Devices, Sunnyvale, Calif.).

To analyze AO concentrations, liposome samples were collected followingAO loading and placed into 7 kDa Mw cutoff dialysis units (Slide-A-LyzerMINI; Pierce, Rockford, Ill.) with isosmotic HEPES as the exchangebuffer. The samples were dialyzed for a total of 4 h with 3 bufferexchanges to ensure complete purification of free AO. Absorbancemeasurements of the purified samples as well as a serial dilution of AOin buffer at λ_(max)=495 nm were taken using a plate reader (SpectraMax;Molecular Devices, Sunnyvale, Calif.) to determine encapsulated AOconcentration. The drug-to-lipid ratio of each resultant sample wasobtained through these absorbance measurements together with theoreticalfinal lipid concentration, as determined by the initial lipidconcentration and given flow rate ratio. The estimate for the finallipid concentration was based on the assumption that all lipids wereincorporated into vesicles in the final solution. In practice, someportion of these lipids may be excluded from the liposomes and remain insolution as small micelles or aggregates, and thus the calculated D/Lvalues reflect conservative estimates for this parameter.

To assess the remote drug loading process following buffer exchange,liposomes were first prepared in a separate microfluidic chip byhydrodynamic flow focusing. Briefly, hybrid PDMS-glass devices with 50μm wide and 300 μm deep microchannels were fabricated, and lipid-ethanolmixture (40 mmol L⁻¹) was injected into the microfluidic device betweentwo sheath flows of ammonium sulfate buffer (250 mmol L⁻¹, pH 4.6). Theflow rate ratio, defined as the ratio of the volumetric flow rate of theaqueous buffer to the flow rate of lipids in ethanol, was set to 20.Total linear flow velocity was set to 12.5 cm s⁻¹, or an equivalentvolumetric flow rate of 112 μL min⁻¹. To reduce vesicle size, themicrofluidic device was operated on a hot plate at 50° C. throughoutsynthesis (Zook, J. M. & Vreeland, W. N. (2010) “Effects oftemperatures, acyl chain length, and flow-rate ratio on liposomeformation and size in a microfluidic hydrodynamic focusing device,” SoftMatter, 6:1352-1360). The resulting liposome size distributions (volumeweighted) were characterized via dynamic light scattering (Nano ZSP,Malvern Instruments Ltd., UK). Volume weighted distributions, i.e. sizesweighted proportionally to their volume, were chosen to represent samplediameters to avoid signal seen by any large aggregates or dust presentwithin the sample.

The liposomes in ammonium sulfate buffer were injected into the inlet ofthe microdialysis chip with isosmotic HEPES (pH 7.6) as the buffercounterflow. AO, an amphipathic dye used as a drug analog for remoteloading experiments, was introduced through a secondary channelimmediately after buffer exchange at a ratio of 1:3 relative to thesample channel volumetric flow rate. Liposome sample velocity was variedfrom 0.17 cm s⁻¹ to 0.44 cm s⁻¹ with AO concentration constant at 0.25mg mL⁻¹ to investigate the effect of flow velocity and residence time onloading concentration and efficiency. AO concentration was varied from0.125 mg mL⁻¹ to 2.5 mg mL⁻¹ (corresponding to D/L values of 0.22 to4.35, respectively) with the flow velocity held constant at 0.26 cm s⁻¹to demonstrate the effect of AO concentration on loading efficiency andmaximum D/L levels in the resulting drug-laden liposomes.

In-Line Liposome Synthesis and Drug Loading

An integrated device containing a liposome formation region,microdialysis buffer-exchange region, and drug-loading region was usedto evaluate the overall process. For liposome formation, lipid-ethanolsolution (20 mmol L⁻¹) was injected into the flow-focusing elementbetween two sheath flows of aqueous ammonium sulfate buffer (250 mmolL⁻¹, pH 4.6). The total volumetric flow rate was 6 μL min⁻¹(corresponding to 0.26 cm s⁻¹ in the dialysis region) with a flow rateratio of 10. The microdialysis counterflow buffer flow rate was matchedto the primary flow rate to minimize the average pressure gradientacross the RC membrane. The drug:liposome sample flow rate ratio was 1:3for all experiments. For drug loading, both AO (0.5 mg mL⁻¹ and 1.0 mgmL⁻¹, corresponding to initial D/L values of 0.22 and 0.44,respectively) and DOX (1.4 mg mL⁻¹, corresponding to an initial D/L of0.44) were investigated. Multiple samples were collected (n=3) for eachtest.

To ensure that free drug remaining in the collection buffer followingremote loading did not affect concentration measurements, collectedliposome samples were further dialyzed off chip with isosmotic HEPES asthe exchange buffer. Two aliquots of each sample, one off-chip and onetwo-fold dilution, were dialyzed for 4 h with 3 buffer exchanges forcomplete purification of free DOX or AO. Absorbance measurements of thepurified samples as well as a serial dilution of DOX or AO in bufferwere compared with standard curves to determine final encapsulated drugconcentration. Size distributions of collected samples were furthercharacterized by dynamic light scattering (Nano ZSP, MalvernInstruments).

Discussion

Counterflow microdialysis provides an efficient method for on-chipbuffer exchange, enabling rapid transport and removal of free ions fromthe sample buffer while preventing loss of nanoparticles during theexchange of small ions across the microdialysis membrane. To evaluateperformance of this approach for establishing transmembrane iongradients in a rapid flow-through format, a PDMS-RC device consistingsolely of the counterflow microdialysis zone was fabricated.Characterization of the device for ion exchange was first evaluated byintroducing buffers at different pH values through the sample andcounterflow ports. Rapid pH change of the sample flow was observed, asrevealed through pyranine fluorescence measurements, with the level ofpH shift roughly proportional to residence time as determined by theapplied flow rate (FIG. 9). As expected, microdialysis performance wasalso found to be dependent on counterflow buffer pH, with a greaterdifference in pH between the sample and counterflow buffers resulting ina larger pH shift at the sample buffer outlet. For the experimentalconditions tested, a maximum shift of 3 pH units was achieved using pH9.6 counterflow buffer and a residence time within the dialysis channelof 83 s. This is substantially faster than bulk scale microdialysis,which can take hours for complete buffer exchange to occur.

A prediction for the transport of ammonium sulfate ions duringmicrodialysis can be made by considering simple diffusion within thesystem. Using a value of 8.0×10⁻⁶ cm² s⁻¹ for the diffusion coefficientof ammonium sulfate in water at room temperature (Leaist, D. G. & Hao,L. (1992) J. Solution Chem., 21:345-350), the diffusion time for acharacteristic length scale given by the microchannel height (37 μm) is1.71 s, significantly smaller than the residence times explored in thiswork which ranged from 42 s to 83 s. To verify this prediction, ammoniumion transport was evaluated through a two dimensional numericalsimulation of the device. The model indicates that the extra-liposomalammonium ion content is reduced by more than 100 times from the initialconcentration sequestered within the vesicles (FIG. 10, panel a, andpanel b), a desired condition for effective remote loading.

A potential issue with the continuous flow microdialysis element ispotential alteration of liposome size during dialysis. Prior to bufferexchange, the microfluidic-synthesized liposomes were found to be 80.8nm in diameter with a very low polydispersity index (PDI) of 0.049. Sizedistributions measured before and after on-chip microdialysis revealedonly a slight increase in mean vesicle size to 91.5 nm, confirming thatthe on-chip counterflow microdialysis element did not significantlyaffect the liposome size. Buffer counterflow eluent was also collectedand examined via light scattering. No detectable signal was observed,revealing that intact liposomes do not escape the membrane and enter thecounterflow during microdialysis.

As noted above, anthracyclines represent an important class of drugs forliposomal encapsulation. Received by nearly every patient undergoingsystemic cancer chemotherapy, anthracyclines are among the most utilizedand effective antitumor drugs developed to date (Hortobágyi, G. N.(1997) “Anthracyclines in the treatment of cancer. An overview. Drugs,54:1-7). Liposomal forms of anthracyclines can provide increasedefficacy with significantly reduced toxicity (Minotti, G. et al. (2004)“Anthracyclines: molecular advances and pharmacologic developments inantitumor activity and cardiotoxicity.” Pharmacol. Rev., 56(2):185-229)enhancing the overall clinical value of the drugs (Allen, T. M. &Martin, F. J. (2004) “Advantages of liposomal delivery systems foranthracyclines,” Semin. Oncol. 31:5-15). Liposomal encapsulation of theanthracycline doxorubicin (DOX) has proven particularly successful fortreatment of a range of cancers (Barenholz, Y. C. (2012) “Doxil®—thefirst FDA-approved nano-drug: lessons learned.” J. Controlled Release,160:117-134). Accordingly, DOX was targeted as a model drug encapsulantto investigate the potential for continuous flow remote drug loadingusing the disclosed microfluidic process.

Initial testing was performed using preformed liposomes injected intothe counterflow microdialysis chip to form the desired transmembrane iongradient, followed by on-chip introduction of AO as a suitable analog toDOX. AO is an amphipathic weak base with similar properties to DOX, andis known to behave similarly to DOX during remote loading with ammoniumsulfate gradients (Clerc, S. & Barenholz, Y. (1998) “A quantitativemodel for using acridine orange as a transmembrane pH gradient probe,”Anal. Biochem., 259:104-111; Zucker, D. et al. (2009) “Liposome drugsloading efficiency: a working model based on loading conditions anddrug's physiochemical properties,” J. Controlled Release, 139:73-80).FIG. 11, panel a presents the measured final D/L and encapsulationefficiency (EE) for AO-loaded liposomes prepared using an initial D/L of0.44 when varying the sample velocity from 0.18 cm s⁻¹ to 0.45 cm s⁻¹,for a total residence time within the mixing channel ranging from 4.25min to 1.7 min, respectively. While a slight increase in both final D/Land EE was observed with increasing flow rate, overall remote loadingusing the microfluidic approach exhibited little dependence on flowvelocity. In contrast, the initial concentration of drug compoundintroduced following ion exchange had a substantial effect on the finalD/L (FIG. 11, panel b). By increasing the initial D/L level, final D/Lvalues up to 1.3 were achieved. This effect is believed to be due to thesignificantly decreased diffusion lengths with increasing AOconcentration, and thus a greater quantity of AO may be loaded when theinitial concentration is higher. Reported D/L values for liposomalanthracyclines produced via conventional bulk-scale remote loading aretypically below 0.25 (Drummond, D. C. et al. (1999) “Optimizingliposomes for delivery of chemotherapeutic agents to solid tumors,”Pharmacol. Rev., 51:691-743), significantly less than the levelsachieved using the disclosed microfluidic process of the presentinvention.

The higher optimal D/L observed for the microfluidic platform can beexplained by the rapid introduction of AO following buffer exchange,together with the use of a high initial D/L and effective on-chip mixingbetween liposomes and encapsulant for decreased diffusion lengths duringdrug loading in the presence of a stable and steep ion gradient.Encapsulation efficiency of the resulting liposomes was observed toincrease with initial D/L, and then began to diminish for initial D/Lvalues exceeding 2.17. This result is in accordance with studies basedon bulk-scale loading based on longer loading periods (hours to days),which suggest EE peaks at an initial D/L of 0.95 and decreases at higherratios (see Zucker, D. et al. (2009) “Liposome drugs loading efficiency:a working model based on loading conditions and drug's physiochemicalproperties,” J. Controlled Release, 139:73-80). This behavior is due toinsufficient intravesicular loading capacity above some limiting D/Llevel, resulting in a lower EE as drug concentration is furtherincreased. The higher optimal D/L observed for the microfluidic platformis believed to result from the highly efficient formation of atransmembrane ion gradient due to rapid microfluidic buffer exchange,followed by immediate interactions between the vesicles and theamphipathic molecules to be loaded. The steep transmembrane ion gradientachieved through rapid buffer exchange enables higher D/L ratios to beachieved through microfluidic remote loading versus bulk scaleprocesses.

After demonstrating buffer exchange, pH adjustment, and remote loadingof AO into preformed vesicles using the microfluidic approach, liposomesynthesis in-line with microdialysis and remote loading of both AO andDOX was performed using an integrated device combining all of theprocess steps in a single flow-through chip. The resulting liposomeswere first characterized for diameter when AO, DOX or buffer wasalternately injected as the drug loading phase (FIG. 12). Under the flowconditions used in these experiments, the resulting liposomes exhibitedan average diameter of 225.5 nm±44.8 nm for the case of buffer withoutamphipathic encapsulant, while a reduction in liposome size to 190.9nm±43.0 nm for DOX-loaded liposomes and 191.5 nm±33.4 nm for AO-loadedliposomes was observed. The reduction in liposome size of approximately15% after remote loading may reflect a change in morphology to acharacteristic “coffee bean” shape resulting from drug crystallizationwithin the vesicles, leading to altered signals during dynamic lightscattering. Confirmation of drug crystallization was verified throughcryogenic transmission electron microscopy (cryoTEM) imaging. A cryoTEMimage of a DOX-loaded liposome is shown inset in FIG. 12.

A summary of measured final D/L values following AO and DOX loadingwithin the integrated microfluidic device is presented in Table 1 below.Prior results for the case of AO loading using preformed liposomesgenerated in a separate chip are also shown for comparison. An initialD/L of 0.44 was selected for DOX, as this ratio was found to maximizeloading efficiency in initial experiments using AO (FIG. 10). The DOXloaded liposomes generated by the microfluidic device had a final D/L of0.32±0.03 (standard deviation), which exceeds the typical D/L of 0.25 orless achieved by conventional bulk remote loading employing overnightincubation. The total on-chip residence time within the microfluidicdevice was less than 3 min. In addition, initial D/L values of 0.22 and0.44 were used for testing AO loading, resulting in final D/L values of0.06±0.01 and 0.32±0.11, respectively.

TABLE 1 Summary of D/L and EE measurements for AO loaded liposomes(preformed liposomes produced in a separate chip) and DOX and AO loadedliposomes (formed in-line with synthesis in a single integrated chip).Similar results were achieved for DOX and AO loading within theintegrated device. Slightly higher final D/L and EE values were observedwhen remote loading was performed in-line with liposome synthesis. Thein-line remote loading process achieves D/L levels exceeding typicalvalues of 0.25 or below achieved by day-long bulk incubation, but withonly a 3 minute on-chip residence time. Liposome Loading diameter Drug/D/L case (nm) agent (initial) D/L (final) EE (%) Preformed  80.8 ± AO0.22 0.05 ± 0.01 23.5 ± 4.2  17.9 0.44 0.24 ± 0.06 55.9 ± 10.0 In-line191.5 ± AO 0.22 0.06 ± 0.01 26.9 ± 2.2  33.4 0.44 0.32 ± 0.11 69.8 ±18.0 190.9 ± DOX 0.44 0.32 ± 0.03 71.8 ± 4.2  43.0

Encapsulation efficiency using the integrated system was also evaluated.Referring to Table 1 above, the DOX-loaded liposomes yielded an EE ofapproximately 72%, lower than typical values for conventional remoteloading which can exceed 99% (Lewrick, F. & Süss, R. (2010) “Remoteloading of anthracyclines into liposomes,” Methods Mol. Biol.,605:139-145). Lower encapsulation efficiency compared to bulk-scaleremote loading is not surprising, since the incubation time is up to 300times lower within the microfluidic system. If desired, EE may beincreased by implementing an additional on-chip region for capturing andrecycling drug following initial remote loading (e.g., such as asecondary dialysis region). Optimizing loading conditions eliminates theneed for off-chip purification, enabling real-time production of stablyencapsulated, highly concentrated liposomal drugs at or near the pointof care.

Direct comparison of in-line system performance with the previousresults from remote loading using preformed liposomes is hampered by thedifferent sizes of each liposome population. When performing in-lineremote loading, fluidic coupling between the upstream liposome formationzone and downstream microdialysis and drug loading zones demands carefuldesign of the channel dimensions, together with appropriate selection ofinlet flow rates. Liposome size and polydispersity are both impacted bythe buffer:lipid flow rate ratio, overall volumetric flow rate, andmicrochannel dimensions selected for effective liposome self-assemblyduring hydrodynamic flow focusing. However, these same parameters alsoaffect microdialysis and remote drug loading performance. While channeldimensions for each functional element in the system can be designedindependently, allowing a degree of decoupling between theseconstraints, conventional microfabrication processes used for devicemanufacture present some limitations. For example, to avoid sagging ofthe microdialysis membrane, a maximum channel width of 1.2 mm was usedfor the buffer exchange zone. Similarly, total volumetric flow rateswere minimized to prevent delamination of the hybrid microfluidic devicedue to excessive internal fluid pressure. As a result of theseconstraints, liposomes formed using the on-line system wereapproximately twice the diameter of their preformed counterparts.Although residual pH gradients decrease with decreasing vesicle size dueto decreased intravesicular volumes, it has been shown that this effectcan be circumvented by including a buffering capacity greater than 300mmol L⁻¹ (Mayer, L. D. et al. (1990) “Characterization of liposomalsystems containing doxorubicin entrapped in response to pH gradient,”Biochim. Biophys. Acta 1025:143-151). Additionally, given thesignificantly larger volume of the in-line liposomes, it is notable thatboth the final D/L and EE values for in-line encapsulation werecomparable to the case of preformed liposome loading, with slightincreases in both values for the larger liposomes. This furtheremphasizes the observation that the initial D/L is an importantparameter of encapsulation performance during remote loading within thecontinuous flow microfluidic system.

Further, liposome concentrations achieved by the microfluidic systemwere typically in the range of 10¹⁰-10¹² liposomes mL⁻¹, depending onexperimental parameters including lipid concentration and flow rateratio. For comparison, reported concentration levels generated bytypical bulk production methods range from 10⁷-10¹² liposomes mL⁻¹(Chen, H. et al. (2010) “Construction of supported lipid membranemodified piezoelectric biosensor for sensitive assay of cholera toxinbased on surface-agglutination of ganglioside-bearing liposomes,” Anal.Chim. Acta, 657(2):204-209; Hitchcock, K. E. (2010) “Ultrasound-enhanceddelivery of targeted echogenic liposomes in a novel ex vivo mouse aortamodel,” J. Controlled Release, 144:288-295), indicating that themicrofluidic technique is an effective and viable alternative forliposomal drug preparation without requiring additional steps to furtherconcentrate the vesicles following on-chip processing.

Thus, by combining liposome synthesis via microfluidic flow focusing,membrane microdialysis for buffer exchange, and in-line introduction ofamphipathic weak bases for remote loading within a single microfluidicdevice, the conventional multi-step bulk-scale processes requiring hoursto days of labor are replaced by the microscale process of the presentinvention, which requires a total on-chip residence time of less than 10minutes, and in some embodiments less than 5 minutes (e.g.,approximately 3 minutes or less).

By taking advantage of the reduced diffusive length scalescharacteristic of microscale flows, the microfluidic method implementsremote loading as a seamless continuous flow process, therebysimplifying yet increasing the robustness of remote loading innanoliposomal drug production. The further ability to perform integratedliposome formation using hydrodynamic flow focusing prior to formationof a transmembrane ion gradient for remote drug loading allows theentire sequence of steps required for liposomal drug production to beperformed as a single in-line and continuous flow process. Themicrofluidic methods and systems demonstrated herein enableexceptionally high drug loading levels, with D/L values above unityeasily achieved, enabling point-of-care production of purified liposomaldrug formulations with minimal drug waste.

HAR-MHF Methods and Comparisons

Microfluidic devices enabling high throughput production of liposomalnanoparticles were demonstrated utilizing high aspect ratio microfluidichydrodynamic flow-focusing (HAR-MHF) methods. HAR-MHF devices featuredmicrochannels having extremely wide channel widths and narrow channelheights. Lipid and buffer microchannels were arranged in multiple layerssuch that the lipid stream was focused into a thin sheet parallel to theplane of the device, unlike prior MHF methods in which the lipid streamis focused perpendicular to the plane of the device. Microfluidicliposome synthesis for large scale applications (e.g. in vivoexperiments, preclinical studies, and point-of-care applications) and atunparalleled production rates was achieved, while retaining the benefitsof effective liposome size control (such as demonstrated using MHF) andreduced levels of polydispersity.

HAR-MHF Device Fabrication

Microfluidic devices were created using a combination of cyclic olefincopolymer (COC) plaques and thin COC films to produce high aspect ratiomicrochannels using simple fabrication techniques without the need forphotolithographic methods or clean room processing. Buffer routingchannels were fabricated in 1 mm thick COC plaques (6013 grade; TopasAdvanced Polymers, Inc., Florence, Ky.) through a precision computernumerical control (CNC) milling machine (MDX-650A; Roland, Lake Forest,Calif.). Thin (50 μm) COC films (6013 grade; Topas Advanced Polymers,Inc.) were used to define a highly uniform channel height throughout theentire width of the mixing channel. The COC films were patterned usingan automated craft cutter (Cameo Digital Craft Cutting Tool, SilhouetteAmerica, Inc., Orem, Utah) to be 5 mm in width, resulting inmicrofluidic devices with an effective 100:1 aspect ratio in thefocusing channel. The resulting intersection between the buffer andlipid phases featured channels which were 100 μm and 50 μm in width(FIG. 13).

Traditional MHF devices were also fabricated using soft lithographictechniques for comparison to the HAR-MHF device, with the flow focusingintersection between the buffer and lipid-ethanol channels designed tobe identical to the HAR-MHF device. Briefly, SU-8 negative photoresist(MicroChem Corp., Newton, Mass.) was spin-coated onto a 4-inch siliconwafer (University Wafer, South Boston, Mass.), exposed to ultravioletlight through a photomask on an automated EVG 620 mask aligner (EVGroup, Germany), and developed to create master molds with raisedfeatures which were used to define microchannel features. The SU-8 moldwas placed in a plastic petri dish and poly(dimethylsiloxane) (PDMS)elastomer (Sylgard 184, Dow Corning Corp. Midland, Mich.) was pouredover the mold. Upon curing, the PDMS was carefully removed from the SU-8mold and inlet and outlet holes were made using a biopsy punch (HarrisUni-Core, Ted Pella, Inc., Redding, Calif.). The bonding surfaces of thePDMS and a glass slide were cleaned using isopropanol and DI water, thenexposed to oxygen plasma in a March Jupiter III Reactive Ion Etcher(Nordson Corp., Concord, Calif.). Final microfluidic device channeldimensions were 50 μm wide and either 25 μm or 250 μm high in the mixingregion, resulting in devices with 0.5:1 and 5:1 aspect ratios (FIG. 13).

Lipid Film and Buffer Preparation

Dimyristoylphosphatidylcholine (DMPC), cholesterol (Avanti Polar LipidsInc., Alabaster, AL) and dihexadecyl phosphate (DCP) (Sigma Aldrich, St.Louis, Mo.) were combined in chloroform (Mallinckrodt Baker Inc.,Phillipsburg, N.J.) at a molar ratio of 5:4:1. The lipid mixture wasprepared in a glass scintillation vial then stored in a vacuumdesiccator for at least 24 h for complete solvent removal. Thedesiccated lipid mixture was re-dissolved in anhydrous ethanol (SigmaAldrich) for a total lipid concentration of 20 mmol L⁻¹ unless otherwisenoted. To assist in visualization during flow focusing experiments, alipophilic membrane dye,1,1′-dioctadecyl-3,3,3′,3′-tetramethylindocarbocyanine perchlorate(DiI-C18; DiI) (Life Technologies, Carlsbad, Calif.) was included intothe lipid mixtures (1 wt %).

Numerical Simulation of Ethanol Concentration Profiles & Fluid Velocity

A computational fluid dynamics simulation was produced to illustrate thevariations in physicochemical properties between the distinctivemicrochannel aspect ratios. The ethanol-water concentration and fluidvelocity profiles of a center stream of ethanol focused by an exteriorsheath of water was represented in a three-dimensional model createdusing COMSOL Multiphysics 4.2 software (COMSOL Inc., Burlington, Mass.).For objective comparison, the flow rate ratio (FRR) in each simulationwas set to 20:1 and the total flow velocity was set to 0.1 m/s(corresponding to volumetric flow rates of 1.5 mL/min, 75 μL/min, and7.5 μL/min for aspect ratios of 100:1, 5:1, and 0.5:1, respectively).

Microfluidic Liposome Synthesis

Liposomes were synthesized through MHF for comparison. A center streamof lipid solvated in ethanol was injected between two streams containingaqueous hydration buffer (FIG. 13, panel a). The FRR, defined as thevolumetric flow rate of the aqueous buffer to that of the solvent, wasset to 10, 15, 20, 30, 40, 50, and 100 for each set of HAR-MHF and MHFexperiments, with linear flow velocity held constant at 0.1 m/s in eachdevice.

In the HAR-MHF device, the FRR was set to 20 with linear flow velocityat 0.1 m/s (corresponding to volumetric flow rates of 1.5 mL/min, 75μL/min, and 7.5 μL/min for aspect ratios of 100:1, 5:1, and 0.5:1,respectively). Liposome populations were characterized through dynamiclight scattering (Nano ZS, Malvern Instruments Ltd., UK).

Results & Discussion: HAR-MHF vs. MHF

Microfluidic processes benefit from high aspect ratio microchannels dueto the diminishing effect of sidewall interaction which presents ano-slip boundary condition and thus disturbs flow profile and chemicalspecies homogeneity (Ismagilov, R. F. et al. (2000) “Experimental andtheoretical scaling laws for transverse diffusive broadening intwo-phase laminar flows in microchannels,” Appl. Phys. Lett., 2000,76:2376-2378; Hertzog, D. E. et al. (2004) “Femtomole mixer formicrosecond kinetic studies of protein folding” Anal. Chem. 76:7169-78).However, microscale features become increasingly difficult to achieve asaspect ratio increases (Ito, H. (2005) “Chemical amplification resistsfor microlithography,” Adv Polym Sci 172:37-245), requiringsophisticated equipment and still failing to exceed values of about 20:1(Becker, H. & Heim, U. (2000) “Hot embossing as a method for thefabrication of polymer high aspect ratio structures,” Sensors ActuatorsA Phys. 83:130-135; Hung, P. J. et al. (2005) “A novel high aspect ratiomicrofluidic design to provide a stable and uniform microenvironment forcell growth in a high throughput mammalian cell culture array,” LabChip, 5:44-8).

HAR-MHF offers a method which achieves unprecedentedly high aspectratios as well as providing the advantages of simplified fabricationmethods used for low aspect ratio features by rotating or reorientingthe axis of diffusion perpendicular to the focusing axis as compared toconfigurations in traditional MHF. Here we demonstrate a HAR-MHF devicefor liposome synthesis within microchannels which have an aspect ratioof 20:1 or greater, more preferably 50:1 or greater, more preferably100:1 or more. Devices having such aspect ratios which would bevirtually impossible to produce through traditional microchannelfabrication techniques. Indeed, devices including microchannels withaspect ratios exceeding 200:1 may be provided in accordance with thedisclosed fabrication techniques, thus providing an alternative methodfor MHF in microchannels with remarkably high aspect ratios.

Populations of liposomes were generated within both HAR-MHF andtraditional MHF devices with varying channel aspect ratios in order todemonstrate how the high aspect ratio affects the resulting liposomepopulations. The microchannel dimensions at the intersection of thelipid and buffer channels and the mixing channel width were identicalfor all 3 devices tested, with the variation being that flow focusingoccurred in the vertical plane within HAR-MHF microchannels whichfeatured an aspect ratio of 100:1, while the focusing axis was in thehorizontal plane in the two traditional MHF devices which featuredmicrochannel aspect ratios of 0.5:1 and 5:1.

Similar to traditional MHF, HAR-MHF produced narrowly distributedpopulations of liposomes in addition to providing the ability to controlliposome size with FRR (FIG. 14). The capacity to inflict major changein liposome size with varying FRR decreased at FRRs above 30 for alldevices, which is in accordance with previous studies of MHF forliposome production (e.g., Jahn, A. et al. (2007) “Microfluidic DirectedSelf-Assembly of Liposomes of Controlled Size,” Langmuir 23:6289-6293;Jahn, A. et al. (2008) “Preparation of nanoparticles by continuous-flowmicrofluidics,” J. Nanoparticle Res. 10:925-934; Jahn, A. et al. (2010)“Microfluidic mixing and the formation of nanoscale lipid vesicles,” ACSnano 4:2077-2087). The various microfluidic systems produced repeatableresults, with modal diameters of the populations of liposomes producedunder a particular FRR remaining consistent (<17% variation) between theHAR-MHF and MHF devices. This comparison study therefore exhibits theability of HAR-MHF to generate nearly monodisperse liposomes of tunablesize as previously demonstrated with traditional MHF.

In addition to modal diameter, the quality of the liposomes producedthrough the various microfluidic flow focusing methods was assessedthrough the comparison of the polydispersity indices (PDIs) of eachpopulation (FIG. 15). Under each FRR tested, the PDI of the populationof liposomes produced through HAR-MHF was lower than the correspondingpopulation produced through traditional MHF, other than FRR 10 in whichHAR-MHF produced a similar value to MHF (5:1). However, lower FRR havepreviously shown higher levels of polydispersity and lower levels ofsize control, so it is not surprising that the PDI was relativelyunchanged despite the varying microchannel aspect ratios.

The lower PDI seen in HAR-MHF-generated vesicles demonstrates that ahigher aspect ratio, and thus more uniform fluid velocity across thewidth of the mixing region, enables the production of more narrowlydistributed populations of liposomes. Moreover, due to the microchanneldimensions, HAR-MHF (aspect ratio 100:1) supported the production ofliposomes at rates 200 times and 20 times faster than the MHF deviceswith aspect ratios of 0.5:1 and 5:1, respectively. Therefore, HAR-MHFpresents a method which enables the generation of microfluidic-enabledliposomes of tunable size with unprecedentedly low levels ofpolydispersity and exceedingly high rates of production as compared totraditional MHF methods.

Three-dimensional numerical simulations further demonstrate the varyingflow conditions within each microfluidic device (FIG. 16). Simulationsof fluid velocity at the midpoint of each channel width was normalizedto distance along the channel height for comparison. As microchannelaspect ratio increases, the effect of the no-slip boundary condition(zero velocity at the channel walls) diminishes, and thus the flowvelocity becomes more homogenous throughout the mixing region andresults in more uniform concentration profiles of the ethanol-lipidthroughout the mixing channel. This feature may explain the observeddecrease in PDI for HAR-MHF versus both MHF microchannel geometrieswhich generate more parabolic velocity profiles.

HAR-MHF: Initial Lipid Concentration

One way to control final liposome concentration, which is important forreaching relevant dose levels for in vivo and preclinical studies,within MHF methods for liposome production is to control the initiallipid concentration. The effect of the initial lipid concentration onthe resulting populations of liposomes produced through HAR-MHF wasexamined by varying the lipid concentration incrementally from 5 mmolL⁻¹ to 80 mmol L⁻¹ (FIG. 17). As demonstrated, increasing the initiallipid concentration (and thus final volumetric fraction of liposomes)did not result in any substantial changes to the resulting liposomepopulations as the concentration varied from 5 mmol L⁻¹ to 40 mmol L⁻¹in terms of both modal diameter (FIG. 17, panel a) or PDI (FIG. 17,panel b). Rather, increasing the initial lipid concentration to 60 mmolL⁻¹ and 80 mmol L⁻¹ caused liposome modal diameter to increase inaddition to resulting in higher PDIs. Therefore, the final liposomeconcentration may be increased by increasing the initial lipidconcentration up to 40 mmol L⁻¹, or 8 times higher than initialtraditional MHF studies, without impinging on the quality of vesiclesproduced.

HAR-MHF: Fluid Velocity

Liposome production rates within MHF techniques may be tweaked somewhatby altering overall flow velocities. However, devices used fortraditional MHF contain microscale channel dimensions which do notsupport high volumetric flow rates due to back pressure. In contrast,HAR-MHF devices contain one dimension which is an order of magnitudehigher than traditional MHF, and therefore back pressure issignificantly reduced and unprecedentedly high flow rates are achieved.

The effect of overall flow velocity on liposome production in HAR-MHFwas assessed (FIG. 18). A decrease in flow velocity from 0.1 m/s to 0.05m/s resulted in an increase in modal diameter (-19% larger) as well asan increase in polydispersity (˜30% increase in PDI). However, anincrease in flow velocity from 0.1 m/s to 0.3 m/s (1.5 mL/min to 4.5mL/min) did not significantly affect liposome quality, with the modaldiameter fluctuations within 5% over these velocities. In addition, thePDI fluctuated slightly across the various flow conditions, but no trendwas observed between flow velocity and polydispersity. Thus, asdemonstrated, liposome production through HAR-MHF is not affected byflow velocity and therefore the primary limitation regarding throughputfrom a single HAR-MHF element is back pressure from the device andtubing inlets.

Comparison of Methods: Liposome Production Rate

An advantage of HAR-MHF is its ability to generate nearly monodispersepopulations of liposomes at extraordinary rates as compared to priormicrofluidic systems. To demonstrate this advantage, typical values forliposome production rates utilizing HAR-MHF were compared to the variousdemonstrated methods of traditional MHF-directed liposome synthesis,including traditional MHF and 3D-MHF (Hood, R. R. et al. (2014) “Afacile route to the synthesis of monodisperse nanoscale liposomes using3D microfluidic hydrodynamic focusing in a concentric capillary array,”Lab Chip 14:2403- 2409), maintaining a constant initial lipidconcentration of 20 mmol L⁻¹ (FIG. 19). HAR-MHF enables the generationof liposomes at ˜100 mg/h lipid, two orders of magnitude higher thantraditional MHF, without sacrificing any of the benefits provided bymicrofluidic liposome synthesis (e.g., controlled size and lowpolydispersity). Similarly, HAR-MHF also produces liposomes nearly twoorders of magnitude faster than the demonstrated capillary-based 3D-MHFmethod, while also providing approximately 10²˜10³ times higher vesicleconcentration than the capillary system due to necessary flow conditionsto achieve particular sizes. The data confirmed that HAR-MHF was abouttwo orders of magnitude faster in microfluidic liposome generation ascompared to prior microfluidic techniques. Further, HAR-MHF systems inaccordance with disclosed embodiments are capable of producingnanoparticles faster than other prior devices utilizing parallelizationfor high-throughput nanoparticle production due to the increased channeldimensions.

Thus, the disclosed systems and data demonstrate HAR-MHF devices areextremely effective for high throughput continuous flow synthesis ofself-assembled nanoscale liposomes. HAR-MHF methods enable theproduction of nearly monodisperse liposomes with tunable diameters, andat unprecedented speeds with further reduced levels of polydispersity ascompared to prior microfluidic systems. HAR-MHF techniques in accordancewith the present invention embody significant implications for drugdelivery applications as it renders microfluidic-synthesized liposomesmore practical for in vivo studies, preclinical trials and point-of-careapplications as compared to other previously demonstrated techniques.

All publications mentioned in this specification are herein incorporatedby reference to the same extent as if each individual publication wasspecifically and individually indicated to be incorporated by referencein its entirety. While the invention has been described in connectionwith specific embodiments thereof, it will be understood that it iscapable of further modifications and this application is intended tocover any variations, uses, or adaptations of the invention following,in general, the principles of the invention and including suchdepartures from the present disclosure as come within known or customarypractice within the art to which the invention pertains and as may beapplied to the essential features hereinbefore set forth.

What is claimed is:
 1. A microfluidic device for synthesis of liposomes,comprising: a sample flow channel; a first inlet channel in fluidcommunication with said sample flow channel; and second and third inletchannels in fluid communication with said sample flow channel, saidfirst, second and third inlet channels converging at a flow focusingregion within said sample flow channel, wherein said sample flow channelhas an aspect ratio (height:width) exceeding 20:1 at said flow focusingregion.
 2. The microfluidic device of claim 1, wherein said aspect ratiois 50:1 or greater.
 3. The microfluidic device of claim 1, wherein saidaspect ratio is 100:1 or greater.
 4. The microfluidic device of claim 1,wherein said first, second and third inlet channels are verticallyoriented relative to each other.